WO2006079797A2

WO2006079797A2 - Apparatus for measurement of analyte concentration

Info

Publication number: WO2006079797A2
Application number: PCT/GB2006/000243
Authority: WO
Inventors: Dawood Parker; John Edwin Enderby
Original assignee: Melys Diagnostics Limited
Priority date: 2005-01-28
Filing date: 2006-01-25
Publication date: 2006-08-03
Also published as: WO2006079797A3; GB0501826D0

Abstract

Apparatus for measuring an analyte such as glucose comprises heated ribbon as an IR source (1) outputting IR light in two opposed directions, an ATR waveguide (4) , having a surface for forming an interface with a body region such as skin, hollow waveguide collimators (2) and (5) , and detectors (3) and (6) for each for measurement and reference wavelength bands outputting an optical absorbance d for the measurement band and an optical absorbance c for the reference band, and producing a time dependant output of the concentration g of said analyte calculated as where A and b/a are constants for a given set of conditions of use, and b and a corresponding to measures d and c respectively, but are determined with water rather than said body region in contact with said optical waveguide surface. Collimator (5) and detector (6) view the waveguide (4) at an adjustable angle α. The effect of temperature on the measurement is compensated for by a reading obtained from a temperature sensor (8) . A final value for g is obtained by applying a predictive algorithm to measurements taken at time intervals .

Description

APPARATUS FOR MEASUREMENT OF ANALYTE CONCENTRATION

This invention relates to apparatus for the measurement of the concentration of an analyte, especially but not exclusively an analyte in a living body.

Apparatus according to the invention may be used in the measurement or monitoring of glucose in fluid, for example body fluid, using optical techniques .

Apparatus according to certain embodiments of the invention is particularly suitable for use in situations in which glucose levels must be closely monitored and/or where glucose measurements must be taken repeatedly, such as in diabetes management .

Diabetes mellitus , abbreviated to diabetes , is the name for a group of chronic diseases that affect the way the body uses food to make the energy necessary for life . Primarily, diabetes is a disruption of carbohydrate metabolism and also affects fats and proteins . The blood glucose level of a person with diabetes may vary considerably, from 40 mM (720 mgdl^"1) to as low as 2 mM ( 36 mgdl^"1 ) . In comparison, the blood glucose level of a person without diabetes varies very little , remaining between 4 mM and 7 mM.

Both insulin-dependent diabetes ( IDDM) and non-insulin- dependent diabetes (MIDDM) are associated with serious tissue complications which typically develop after ten to twenty years ' duration of the disease .

In the management of diabetes , the regular measurement of glucose in the blood is essential in order to ensure correct insulin dosing . Furthermore , it has been demonstrated that in the long term care of the diabetic patient better control of the blood glucose levels can delay, if not prevent , the onset of retinopathy, circulatory problems and other degenerative diseases often associated with diabetes . Thus there is a need for reliable and accurate self-monitoring of blood glucose levels by diabetic patients . Since the late 1970s an increasing number of diabetic patients , mostly those suffering from IDDM, have been measuring their own blood glucose concentrations using finger-prick capillary blood samples . Self blood glucose monitoring is used by diabetics in the home to detect hypoglycaemia or hyperglycaemia and therefore to determine the corrective action required, such as taking extra food to raise the blood glucose concentration or extra insulin to lower the blood glucose concentration . These measurements , which are made using a low- cost , hand-held blood glucose monitor, also allow the physician to adjust the insulin dosage at appropriate times so as to maintain near normoglycaemia .

These blood glucose monitors use either reflective photometry or an electrochemical method to measure blood glucose . concentration . The finger or earlobe of the patient is pricked with a sterile lancet and a small sample of blood is placed on the test strip . After analysis , the monitor displays the blood glucose concentration.

The main disadvantages of self blood glucose monitoring systems are : poor patient acceptance because the technique is painful ; only intermittent assessment of diabetic control is possible ; readings during the night or when the patient is otherwise occupied are not possible .

It is estimated that less than half of the IDDM patients in the US perform self blood glucose monitoring . However, the US National Institutes of Health has recommended that blood glucose testing should be carried out at least four times a day, a recommendation that has been endorsed by the American Diabetes Association. This increase in the frequency of blood glucose testing imposes a considerable burden on the diabetic patient , both in financial terms and in terms of pain and discomfort , particularly in the long-term diabetic who has to make regular use of a lancet to draw blood from the fingertips . Thus , there is clearly a need for a better long-term glucose monitoring system that does not involve drawing blood from the patient . There have been a number of proposals for glucose measurement techniques that do not require blood to be withdrawn from the patient . Many of these methods measure blood glucose concentration using near infrared (NIR) spectroscopy to analyse the glucose concentration in the blood vessels in the skin . For example, US 5 , 028 , 787 and 5 , 362 , 966 (Rosenthal et al . ) describe devices for measuring glucose concentration in a fingertip on passing NIR ( 600 to 1100 ran) light through the finger to the detector, and a similar device that measures the glucose concentration in a vein by shining NIR light on to one part of the vein in the wrist and detecting the light emerging from another part of the vein nearer the elbow. Disadvantages of the use of NIR spectroscopy are that the measurement suffers from interference by other optical absorbers in the tissue and is also dependent on blood flow in the skin . The absorbance peaks of water at these wavelengths are variable in position with temperature such that shifts in water absorbance induced by temperature changes of less than a degree are greater than those induced by a 20 mg/dl variation in glucose concentration . Accordingly, US5 , 362 , 966 describes the inclusion of temperature measuring means and the use of the temperature measurements in adjusting the measured value of the glucose concentration.

It has been observed that the concentration of analytes in subcutaneous fluid correlates with the concentration of those analytes in the blood . There have been several techniques developed for measuring the glucose concentration in the tissue ; examples include subcutaneous implants containing glucose binding assays , such as that described in W091/09312 , in which the degree of binding of the assay component to glucose is indicated by the intensity of fluorescence produced by the assay. This type of device is minimally invasive , as the fluorescence signal is able to be read remotely and so the extraction of a blood sample is avoided, but it does requires an invasive operation to install the implant . Therefore , it is apparent that there is a need for an improved method of glucose concentration measurement that may be used regularly by a diabetic patient in the home, which is cheap and simple to use , as well as being accurate and non-invasive in nature . It is an aim of the present invention to provide such a device .

Invasive glucose measurement techniques , however, may still be widely used. Continuous monitoring of the concentration of analytes in the body of a patient undergoing critical care in a hospital or the like is essential , particularly the monitoring of glucose concentration . Currently, as for the self-monitoring of blood glucose , the usual technique involves drawing blood from a patient using a sterile lancet and analysing that sample using reflective photometry or an electrochemical method. This has the disadvantages that the glucose concentration may be measured only intermittently, in addition to the discomfort caused to the patient . There is , therefore, a need for a device which can provide accurate real-time concentration measurements of tissue glucose levels . The use of an invasive device allows readings to be taken as frequently as necessary, and also requires only invasive operations to install and remove the device to be performed on the patient, instead of repeated drawing of blood .

Another aim of the present invention is therefore to provide an invasive device for continuous measurement of levels of analytes , particularly glucose , in the living body.

It is known that an optical waveguide, such as a plate, prism, or optical fibre may pass light along its length by the process of total internal reflection. In the case of a prism, total internal reflection occurs when the angle of incidence of the light entering the prism exceeds the critical angle, θ_c :

where n₂ is the refractive index of the region outside the prism, and n_± is the refractive index of the prism. At each reflection within the prism, an evanescent wave penetrates a small distance into the surrounding medium. It is therefore possible to obtain an absorption spectrum for the regions of the surrounding medium in contact with the outer surface of the prism. The same is true where the waveguide takes another form, such as a fibre; this technique is known as attenuated total reflectance (ATR) spectroscopy. The signal-to-noise ratio for this spectrum increases with the number of reflections within the waveguide achieved with evanescent wave interaction with the absorbent before detection of the emerging light . The use of a material with a high refractive index for the prism is therefore preferred, as this minimizes the critical angle θ_c and thereby maximizes the possible number of reflections within the waveguide . It is estimated that the evanescent wave generated from the reflections of mid infrared (mid-IR) light in the wavelength range 5-10 (Jm in an ATR prism placed against the skin surface of a human fingertip will penetrate to a depth of around 10-30 μm, the^' penetration depth being dependent on the wavelength of light used and the angle, of incidence . This depth is just below the stratum corneum, and therefore the light penetrates the region in which interstitial fluid is present . Interstitial fluid is known to have a glucose concentration that correlates with the blood glucose concentration, and so absorbance data from the interstitial fluid in an appropriate wavelength range may be used to calculate a blood glucose concentration.

The use of ATR methods in non-invasive devices for the measurement of glucose concentration is known . For example, US2003 /0031597 ( Sota et al . ) discloses a method of using an ATR prism to collect a mid-IR spectrum of an obj ect containing glucose placed on the outside of the prism, and comparing that spectrum with a library of spectra taken at different glucose concentrations in order to produce a measurement of the glucose concentration in the obj ect . Such methods have not proved reliable in practice and require undesirable computing complexity and the acquisition of numerous reference spectra tailored to a specific individual .

Japanese application no . 11-363967 (Yoshikawa Osamu) shows an apparatus for providing infrared light of a suitable wavelength for glucose absorption to a diamond ATR prism placed on a surface, and a detector for the emerging light, in order to measure the glucose concentration in a sample .

Japanese patent application no . 09-344435 discloses a device for measuring glucose concentration in a living body by placing an ATR crystal , whose surface is masked so the sample area remains constant , against a body part , supplying infrared light to the crystal , and measuring the absorbance of the emerging light .

WO01/79818 and US6421548 (Berman et al . ) disclose a device incorporating an ATR prism to be placed against the skin surface and held there by a pressure device . Mid-IR light is focused into the prism, and the. emerging light is split and focused to two detectors which measure the absorbance at two wavelengths , one at which the concentration of glucose affects the absorbance strongly, and a reference wavelength in the range 8.25 to 8.75 μm, whose absorbance is said to Ije independent of the glucose concentration . Subtraction of the latter absorbance value from the former absorbance value gives a result which may be correlated with a glucose concentration . The absorbance spectrum of glucose over a wavelength range of 5 to 12 μm is shown in Figure 2 in WO01/79818. However , although Figure 2 of WO01/79818 gives the impression that the absorbance in the range of 8.25 to 8.75 μm is independent of glucose concentration, this is in fact not the case . Also, adjacent this region on each side in the spectrum are strongly absorbing areas , so that if the cut off of the filters used to define the reference wavelength band is not perfect , even more substantial glucose influence on the absorbance of the reference wavelengths will occur . Further , the glucose concentration is calculated as C_x = Jc(A₂ - A₁ )

where k is a constant determinable by calibration and A₂

and A₁ are ratios corresponding to explained

below respectively . In fact , as we explain herein, k in this equation is not adequately represented by a constant . This arises because the absorbance of the measurement wavelength band and the reference wavelength band are not equal , the absorbance at both wavelength bands by the sample are time dependent and depend on the degree of hydration of the stratum corneum and the effective path length in the sample depends on non-glucose influences on the refractive index of the tissue ( including its hydration) , the contact area and other uncontrolled variables . It is desirable to improve on these previous proposals , especially in the context of invasive and non-invasive measurements of glucose .

Several different aspects of the present invention are described herein. They may all be used in combination as specifically exemplified -below, or any one or more of them may be combined without use of the others .

In a first aspect , the invention provides apparatus for measuring analyte concentration in a living body, comprising a source of optical radiation, an optical waveguide having a surface for forming an interface with a body region of said living body, and at least a first detector system, comprising at least one detector , for quantitative detection of at least a measurement wavelength band of said radiation for which an absorption characteristic of said analyte is dependent on the analyte concentration adj acent said optical waveguide surface in contact with said body region to produce an optical absorbance measure d and for quantitative detection of at least a reference wavelength band of said radiation to produce an optical absorbance measure c, wherein said source of optical radiation is positioned to supply said radiation to said optical waveguide to interact as an evanescent wave with the body region containing the analyte via the interface between the body region and the optical waveguide, and the or each detector of the first detector system is positioned to receive said radiation from said optical waveguide after said interaction, and wherein the reference wavelength band is selected to be in a wavelength range for which an absorption characteristic of said analyte is substantially less dependent on the analyte concentration, and computation means for receiving and processing an electronic output provided by at leas t said first detector system representing said quantitative detection of said measurement and reference wavelengths of said radiation to produce an output indicative of the concentration of said analyte, wherein said output is calculated as being directly proportional to an arithmetical combination of said quantitative measures d and c and a factor h/a equivalent to

wherein g is the analyte concentration, A is constant for a given set of conditions of use, F(c) is a function of c, and h/a is constant for a given set of conditions of use .

The computation means may operate such that said function

F (c) is so that g is calculated as an arithmetical

combination of said quantitative measures d and c and a factor h/a. equivalent to

It is apparent that in contrast with the formula used by Beritian et al , the analyte concentration is calculated according to the first aspect of the invention ( and preferably also in the practice of the other aspects described below) taking account of sample dependent alterations in the value of Berman' s ' constant ' k.

Where the analyte is glucose, this form of function is appropriate for analyte concentrations over the whole of the range of physiological interest and will generally produce a result within ¹A itiM/1 of that obtained from a finger prick and blood analysis .

Of course, one can employ -a different function that produces an equivalent result . Thus , for values of d~c over

about 0 : 45 , we find that is substantially proportional

to , the ratio between the two being apparatus dependent . d-c

Accordingly, within this concentration area , one can use the equivalent expression :

Said computation means may be configured for receiving and processing an electronic output provided by said detector system representing optical absorbance values b and a for said measurement and reference wavelength bands respectively, and corresponding to measures d and c respectively, but which are determined with an aqueous reference material rather than said body region in contact with said optical waveguide surface . As an alternative to making measurements of the values a and b, said computation means may be configured for receiving utilising a calculated or separately determined value of b/a. Said calculated or separately determined value of b/a, represents an estimate of the ratio of optical absorbance values b and a which would be produced for said measurement and reference wavelength bands respectively if the apparatus were to be operated with an aqueous reference material rather than said body region in contact with said optical waveguide surface . Preferably, said aqueous reference material is water.

Said detector system preferably comprises a measurement optical filter selectively passing said measurement wavelength band and a reference optical filter selectively passing said reference wavelength band .

For use in each aspect of the invention, preferably said first detector system comprises respective measurement and reference detectors positioned to receive the filtered output of said measurement and reference optical filters respectively. If a single detector is used, it would be possible to apply different filtering at different times and to derive the measurements at closely spaced intervals , but with additional complications .

Preferably, the apparatus comprises a second detector system positioned to receive the mid-IR radiation output from said source without passage through said optical waveguide . Said second detector system may comprise at least one detector, for quantitative detection of at least said measurement wavelength band of said radiation and for quantitative detection of at least said reference wavelength band of said radiation . Preferably, said second detector system comprises a measurement optical filter selectively passing said measurement wavelength band and a reference optical filter selectively passing said reference wavelength band and preferably, said second detector system comprises respective measurement and reference detectors positioned to receive the filtered output of said measurement and reference optical filters respectively.

This enables one to derive measurements ( J) of the intensity of the radiation at each of the measurement (λ_raeas) and reference (λ_ref) wavelengths for light that has passed through the interface region ( I_N) and for light that has passed directly from the source to a detector ( J₀) . In principle, the intensity of the light generated could be determined other than by measurement in this way, e . g . from the power applied to the light source .

Thus , suitably:

a — for water, or more generally the aqueous

reference material ;

for water, or more generally for the

aqueous reference material ;

for the body region; and

for the body region .

The computation means may be a suitably programmed or configured digital or analogue computer .

Means may be provided for accepting input values of one separately determined analyte concentration g as a calibration value, and for establishing the value of A at a known temperature , pressure between the surface and the body region at the interface and angle of incidence of the light on the interface . Alternatively, at least in the context of measuring glucose as the analyte with skin being applied to said surface at said interface, a value of A in the range 4±2 may be used, without calibration of the device using a separately-determined value of g. This range of values of A is observed empirically to be the range into which the maj ority of patients ' A value will fall . As there is no clinical need for a highly accurate measure of glucose concentration, use of a value in this range will produce a glucose concentration indicative of whether the patient ' s current glucose level is normal , too high or too low. However , if a more accurate glucose concentration determination is required ( say to within about l/2mM/l ) , it is preferred to use a calibration step to find a value of A specific to the patient , as described above . J₀ is in practice preferably measured by the second detector system of the preferred forms of the apparatus of the invention, i . e . is radiation from the back face of the resistance heating element used as the preferred radiation source . The measured value will bear a proportional relationship to the true value of J₀.

Especially where the analyte concentration actually measured is not the same as but correlates with an analyte concentration of specific interest (e . g . where the measured concentration is that of glucose in interstitial fluid, but that of specific interest is that in the blood) , the measured analyte concentration may be converted to an indicated analyte level in terms of the concentration of specific interest , suitably by the computational device mentioned above .

The invention according to the first aspect of the invention includes a method of measuring an analyte concentration using apparatus as described above . Optionally, such a method may include measuring the ratio Jb/a.

Generally, the optical waveguide will be in the form of a plate (or prism) or in the form of an optical fibre . Preferably, the optical waveguide conducts said radiation by . -

-13-

multiple total internal reflections at said surface providing said interface such that light is reflected at said surface at least about 25 times in passing through said waveguide . Where the optical waveguide is in the form of a plate or prism on which a finger or fingers are to be placed, the number of reflections obtained within the sampling length will depend on the refractive index and the thickness of the plate or prism.

Preferably, the thickness is not more than 3 mm, suitably 2 mm . Preferably, the measurement wavelength band is in the mid-IR region and the reference wavelength band is in the mid- IR, near-IR, or visible region.

Where the analyte is glucose, for the reasons outlined ^■ below, the measurement wavelength band is preferably in the range 6.6 to 14.0 μm or is in the range 8.8 to 10.5 μm .

In a second aspect , the invention provides apparatus for measuring analyte concentration in a living body,, comprising a source of at least mid-IR radiation, an optical waveguide having a surface for forming an interface with a body region of said living body, and a detector system, comprising at least one detector, for quantitative detection of at least a measurement wavelength band of said radiation for which an absorption characteristic of said analyte is dependent on the analyte concentration adjacent said interface and for quantitative detection of at least a reference wavelength band of said radiation, wherein said source of mid-IR radiation is positioned to supply said radiation to said optical waveguide to interact as an evanescent wave with the body region containing the analyte via the interface between the body region and the optical waveguide, and the or each detector of the detector system is positioned to receive said radiation from said optical waveguide after said interaction, and wherein the reference wavelength band is selected to be in a wavelength range for which an absorption characteristic of said analyte is substantially independent of the analyte concentration . This is in contradistinction to the apparatus of Berman et al described above in which the reference wavelength band is not in a region where absorbance is independent of the analyte (glucose) concentration . The apparatus may further comprise a computational device such as a digital or analogue computer for receiving and processing an electronic output provided by said detector system representing said quantitative detection of said measurement and reference wavelengths of said radiation to produce an output indicative of the concentration of the analyte . As described below, the computer may also deal with compensation of the measured values for ambient temperature at the interface .

The form of the waveguide and the choice of measurement and reference wavelengths may all be as described in accordance with the first aspect of the invention .

There are two criteria which desirably are to be satisfied in choosing the reference wavelength:

(a) Commercial standard filters have a combined half-width and central wavelength (CWL) positional uncertainty of +/- 0.3 |im. Thus the choice of the CWL should be such that there is no glucose contribution to the spectrum in the region CWL +/- 0.3 μm.

(b) The spectrum of the stratum corneum should be free of absorption bands at the chosen wavelengths , otherwise there may be a significant temperature effect .

In addition, it is desirable that the water spectrum is free from strong absorption bands in the range CWL +/- 0.3 Um. The reference wavelength is preferably in the range 5.0 to 5.6 Um, which satisfies all of the above . Alternatively, the reference wavelength may be in the range 1.8 to 2.6 μm, which satisfies criteria ( a) and (b) above . In these regions , the absorption is substantially less dependent on the glucose concentration than in the 8.5 μm region suggested by Berman. This region does not satisfy criterion ( a) above, as at 8.2 and 8.8 μm, the ratios of the reference to the measuring channel are 10% and 30% respectively. For the reference region at 5.3 μm, the same ratio is less than 0.1% .

The width of the wavelength bands used may be so small that the radiation is effectively monochromatic or may be up to the whole of the ranges specified above .

Preferably therefore, the absorption coefficient of the analyte at the reference wavelength band is no greater than 1% when measured on a solution of the analyte in a suitable solvent , which where glucose is the analyte is suitably water .

Alternatively, the ratio of the absorption coefficient of the analyte at the measurement wavelength band to the absorption coefficient of the analyte at the reference wavelength band is preferably greater than 99% when measured on a solution of the analyte ( e . g . glucose) in a suitable solvent (e . g . water) .

The apparatus may provides an output measure of analyte concentration derived from the ratio

where as indicated above I_N is the intensity of said radiation after interaction with the body region containing the analyte at said interface and J₀ is the intensity of the radiation in the absence of said interaction at each of the measurement and reference wavelength bands , indicated by _meas and _ref respectively . This ratio may of course be calculated by the computer referred to above .

Means may be provided for accepting input values of at least two separately determined analyte concentrations Sr₁ and g_∑ as calibration values and for solving for calibration values of constants D and E in equations R₁ = D + E*g_x and

R₂ = D + E*g₂ where R₁ is the measured value of R at analyte concentration g_x and JR₂ is the measured value of R at analyte concentration g_2, and for applying the calibration values of D and E for determining said concentration g in further analyte concentration determinations from measured values of R.

More preferably, said computation means operates , as in the first aspect of the invention, so that the output value is calculated as being directly proportional to an arithmetical combination of said quantitative measures d and c and a factor b/a equivalent to

wherein g is the analyte concentration, A is constant for a given set of conditions of use, F (c) is a function of c, and b/a is constant for a given set of conditions of use .

As before the computation means may operate such that

1 said function F (dc) is so that g is calculated as an

arithmetical combination of said quantitative measures d and c and a factor b/a equivalent to the equation

where, as before :

for water, or more generally for the

aqueous reference material ; for water, or more generally for the

aqueous reference material ;

for the body region; and

for the body region.

Means may be provided for accepting input values of one separately determined analyte concentration g as a calibration value, and for establishing the value of A at a known temperature , pressure between the surface and the body region at the interface and angle of incidence of the light on the interface .

Alternatively, as in connection with the first aspect of the invention, at least for measuring glucose as analyte via the skin, a value of A in the range 4+2 may be used, without calibration of the device using a separately-determined value of g. This range of values of A is observed empirically to be the range into which the maj ority of patients ' A value will fall . As there is often no clinical need for a highly accurate measure of glucose concentration, use of a value in this range will produce a glucose concentration indicative of whether the patient ' s current glucose level is normal , too high or too low.

However, if a more accurate glucose concentration determination is required, it is preferred to use a calibration step to find a value of A specific to the patient , as described above .

Features indicated to be preferred in connection with the first aspect of the invention above are also preferred for use in relation to the second aspect of the invention and the two aspects of the invention may be employed together .

The second aspect of the invention includes a method of determining an analyte concentration by the use of the described apparatus . In a third aspect, the invention provides apparatus for measuring analyte concentration in a sample, comprising a source of electromagnetic radiation emitting said radiation in at least a first direction and a second direction suitably opposite to said first direction, an optical waveguide having a surface for providing an interface with a sample placed in contact therewith and positioned to receive said electromagnetic radiation emitted in said first direction to interact as an evanescent wave with a said sample via said interface between the sample and the optical waveguide surface , a first detector system for detecting said electromagnetic radiation emitted in said first direction after passage through said waveguide, a second detector system for detecting said electromagnetic radiation emitted in said second direction, each said detector system providing an electronic output , and computational circuitry, e . g . a computer, for receiving and processing said electronic outputs to derive therefrom the concentration of the analyte .

More generally, this aspect of the invention includes optical apparatus having a source of optical radiation and having first and second optical paths therefrom to respective detectors , wherein said source is a resistance heating element and said paths originate at said source and diverge therefrom, preferably at an angle of greater than 10 degrees , more preferably greater than 60°, e . g . are in opposite directions .

The sample in this aspect of the invention may be a body region as before but may simply be a liquid or solid sample placed in contact with the waveguide surface for in vitro measurement . This aspect of the invention is particularly relevant to the measurement of a large range of analytes .

The analyte concentration may be derived from the ratio of the outputs of the first and second detector systems , preferably by taking the logarithm of said ratio and more preferably by taking the logarithms of such ratios measured at different wavelengths and establishing the ratio between said logarithms of ratios .

As in previous aspects of the invention, said first detector system may comprise a measurement optical filter selectively passing a measurement wavelength band and a reference optical filter selectively passing a reference wavelength band and may comprise respective measurement and reference detectors positioned to receive the filtered output of said measurement and reference optical filters respectively. Similarly, said second detector system may comprise at least one detector, for quantitative detection of at least said measurement wavelength band of said radiation and for quantitative detection of at least said reference wavelength band of said radiation, and said second detector system may comprise a measurement optical filter selectively passing said measurement wavelength band and a reference optical filter selectively passing said reference wavelength band. Also , said second detector system may comprise respective measurement and reference detectors positioned to receive the filtered output of said measurement and reference optical filters respectively.

As in the previous aspect of the invention, the apparatus may provide an output measure of analyte concentration derived from the ratio

Similarly, the apparatus may comprise means for accepting input values of at least two separately determined analyte concentrations g_x and g_∑ as calibration values and for solving for calibration values of constants D and E in equations R₁ = D + E*g_x and R₂ = D + E*g₂ where R₁ is the measured value of R at analyte concentration g^ and R₂ is the measured value of R at analyte concentration g₂, and for applying the calibration values of D and E for determining said concentration g in further analyte concentration determinations from measured values of said ratio R.

More preferably, the apparatus provides an output measure of analyte concentration g derived as described in relation to the first aspect of the invention from the equation

g = A*F(d,c)*

Similarly, means may be provided for accepting input values of one separately determined analyte concentration g as a calibration value, and for establishing the value of A at a known temperature, pressure and angle .

Alternatively, a value of A in the range 4+2 may be used, without calibration of the device using a separately-determined value of g.

Features indicated to be preferred in connection with the first or second aspects of the invention above are also preferred for use in relation to the third aspect of the invention and any two aspects of the invention may be employed together .

Preferably, the detector incorporates filters for the measurement and reference wavelengths of interest . Preferably, the detector incorporates a filter for each available channel .

As discussed above, temperature compensation has been found essential in near-IR ATR devices for measuring glucose because of the temperature-induced shift in water absorption bands in this wavelength region . Such considerations do not apply in the mid-IR regions of interest , due to the absence of water absorption bands close to 5.3 μm or 7-14 μtα. However , we have found that compensation for temperature is still desirable when working in the mid-IR.

Thus , it has been found that the measurement of glucose concentration in the interstitial fluid is dependent on the temperature of the body region containing the glucose because the absorption coefficient of glucose itself varies with temperature when glucose is present in this environment . It is therefore desirable to correct the measured value of glucose concentration to obtain a standardized value which may be compared with previous values . This is of particular significance in diabetics as the disease is known to cause poor circulation in the extremities , and if the glucose concentration measurement is taken from the fingertips the variation in ambient temperature may be important . Although the IR spectrum of water is relatively featureless in the measuring region, the IR spectrum of the stratum corneum is not . Thus , there will be temperature effects due to band shifts (as described above for water in the near-IR) which are likely to be strongly subj ect-dependent and must therefore be measured and corrected for .

Accordingly, the invention provides in a third aspect apparatus for measuring glucose concentration in a living body, comprising a source of at least mid-IR radiation, an optical " waveguide having a surface for forming an interface with a body region of said living body, a temperature measuring device measuring or estimating the ambient temperature of the interface , and a detector system, comprising at least one detector, for quantitative detection of at least a measurement wavelength band of said radiation for which an absorption characteristic of glucose is dependent on the glucose concentration adjacent said interface and for quantitative detection of at least a reference wavelength band of said radiation, wherein said source of mid-IR radiation is positioned to supply said radiation to said optical waveguide to interact as an evanescent wave with the body region containing the glucose via the interface between the body region and the optical waveguide , and the or each detector of the detector system is positioned to receive said radiation from said optical waveguide after said interaction, the detector system and the temperature measuring device providing electrical inputs to computational circuitry, e . g . a computer, adapted to compute a glucose concentration compensated for the ambient temperature .

Preferably, when using the absorbance ratio method of equation ( 1 ) , said compensation comprises characterising the absorbance ratio change with temperature using non-linear regression, and, by extrapolating from these characteristics , obtaining the absorbance ratio and hence glucose concentration at a chosen reference temperature .

Alternatively, when using the expression in equation ( 2 ) to determine g, the change in the value of A with temperature may be characterised, in order to obtain a value of A corrected for the temperature at which the measurement is taken for use in equation ( 1 ) . Preferably, the change in the value of A with temperature is characterised using the expression: A(T) = A_Q[I- P(T -T₀)] , where : A (T) is the value of A at temperature T °C;

A₀ is the value of A at a reference temperature T₀ °C; p is a constant for a given set of conditions of use and has been found to have a value of approximately 0.2. The third aspect of the invention can be employed in combination with either or both together of the first and second aspects of the invention and each and all of their preferred features .

The third aspect of the invention includes a method of compensating for the effect of temperature in accordance with the operation of the above described apparatus .

It is found that when a skin surface is placed on an ATR interface and the glucose concentration is measured as described above, the reading changes over a period of minutes . Whilst the cause of this is as yet uncertain, it is believe to be due to progressive hydration of the stratum corneum by transport of water into the stratum corneum from the interior, with loss of water by evaporation being temporarily blocked by the occlusive effect of the ATR surface . Such changes in water content of the stratum corneum are described by Cole et al , Proc . SPIE, 4276 p 1-10 , 2001 , 'Terahertz Imaging and Spectroscopy of Human Skin, In-vivo ' .

Whatever the underlying mechanism, we find that the indicated glucose concentration is initially low but rises towards a steady state level in an essentially exponential manner . Accordingly, according to a fourth aspect of the invention, there is provided apparatus for measuring analyte concentration in a living body, comprising a source of optical radiation, an optical waveguide having a surface for forming an interface with a skin region of said living body, and at least a first detector system, comprising at least one detector, for quantitative detection of at least a measurement wavelength band of said radiation for which an absorption characteristic of said analyte is dependent on the analyte concentration adj acent said optical waveguide surface in contact with said skin region to produce an optical absorbance measure d and for quantitative detection of at least a reference wavelength band of said radiation to produce an optical absorbance measure c , wherein said source of optical radiation is positioned to supply said radiation to said optical waveguide to interact as an evanescent wave with the body region containing the analyte via the interface between the skin region and the optical waveguide , and the or each detector of the first detector system is positioned to receive said radiation from said optical waveguide after said interaction, and wherein the reference wavelength band is selected to be in a wavelength range for which an absorption characteristic of said analyte is substantially less dependent on the analyte concentration, and computation means for receiving and processing an electronic output provided by said detector system representing said quantitative detection of said measurement and reference wavelengths of said radiation to produce an output indicative of the concentration of said analyte, wherein said electronic output is dependent on time elapsed from the placing of the skin region on said surface and said computation means operates to predict a maximum value for the measured analyte concentration from values of said output obtained prior to said output reaching a stable level .

Predicted maximum values for the analyte concentration may be repeatedly calculated, e . g . by fitting all the measured concentration values to a curve fitting algorithm, and the rate of change in said predicted maximum values may be determined and a reading of analyte concentration may be output when said rate of change falls below a preset level .

In each aspect of the invention, the source of radiation is preferably an electrical resistance heating element and a source of electrical current connected to pass therethrough .

Suitably, said heating element is of foil , e . g . Nichrome™ foil . The heating element may be caused to produce pulses of radiation by charging and discharging a capacitor through the element . The apparatus may be compact and powered by a rechargeable or replaceable battery .

The or each detector of the or each detector system may preferably be pyroelectric detectors .

Two principal types of use are envisaged for apparatus according to the invention, non-invasive measuring devices and invasive measuring devices . Thus , in the non-invasive type of device , said interface may be positioned and adapted to be contacted by the skin of a subj ect as said body region . To assist the user said surface may be marked to indicate one or more locations on which said subj ect is to place at least one fingertip to constitute said body region for the assay. For instance , the surface may be marked to indicate locations on which to place either both index fingertips or two adjacent fingertips of the patient . Preferably, there are defined areas provided at the interface on which the fingertips are placed, and other areas may be masked so that evanescent waves are only able to pass through these defined areas . For example, the masked areas may be created by depositing a layer of gold on required areas of the interface .

For consistency of results , a pressure-maintaining member is desirably provided to maintain adequate pressure of the fingertips against the said surface . This may take the form of a clamp which is lowered over the fingers on the surface of the waveguide to apply a specific pressure . This helps to ensure that the optical density of the fingertips remains constant during the measurement of the glucose concentration. Alternatively or additionally, the pressure applied to the said surface is measured by a pressure sensor . This may be linked to an audible or visual signal to the user when the pressure applied is within a predetermined range of tolerance .

In previous proposals , optical elements such as lenses and beam splitters have been used to guide IR radiation from a source to an ATR element such as a plate . However , for transparency and suitable -refractive index, these elements have to be fabricated from materials such as zinc selenide and are prohibitively expensive for use in a consumer device . Also, they give rise to excessive optical path lengths . This makes the whole apparatus undesirably large for home use . Also, long path lengths , particularly if they are not identical , may lead to strong, unquantifiable and undesirable water vapour absorbance . According to the various aspects of this invention, the radiation is preferably directed from the source towards the optical waveguide and/or from the optical waveguide to the or a detector system, using a hollow fibre waveguide as a collimator , e . g. a hollow silica collimator . This reduces both the cost and bulk of the apparatus as compared to the use of conventional optics . It is desirable that the apparatus can be fine tuned to optimise measurements according to the character of the skin of a particular subj ect , which can affect the effective refractive index at the interface . For this purpose, it is preferred that said optical waveguide and said detector system are mounted for relative movement such that the angle at which the light exit face of the waveguide plate is viewed by the detector system is variable by relative movement of the detector system and the optical waveguide . Thus , according to a fifth aspect of the invention, there is provided apparatus for measuring analyte concentration in a sample , comprising a source of electromagnetic radiation emitting said radiation in at least a first direction, an optical waveguide having a surface for providing an interface with a sample placed in contact therewith and positioned to receive said electromagnetic radiation emitted in said first direction to interact as an evanescent wave with a said sample via said interface between the sample and the optical waveguide surface at a plurality of different angles of incidence with said interface, at least a first detector system for detecting said electromagnetic radiation emitted in said first direction after passage through said waveguide , wherein the said first detector system is mounted for relative movement with respect to a light exit face of the optical waveguide from which the electromagnetic radiation exits the waveguide after interaction with the sample such that an angle at which the light exit face is viewed by the detector system is variable by relative movement of the detector system and the optical waveguide to thereby select for detection the radiation which has a desired angle of incidence to said interface . This feature may be adopted in the practice of each of the other aspects of the invention .

Preferably, the angle formed between the interface of the ATR crystal or other waveguide and the light path therefrom to the detector may be varied to the extent that at least a 15 ° variation in angle is possible . For use in non-invasive measurement according to the invention, the optical waveguide is preferably an ATR crystal or prism which is preferably a zinc selenide (ZnSe) prism. Suitably, the prism may be capped with CVD diamond, whose refractive index matches that of zinc selenide .

As indicated above, the various aspects of the invention are also applicable to invasive measurements where the body region is an internal tissue of the body. For this purpose, the optical waveguide may comprise an optical fibre having a proximal clad portion and a distal unclad portion for insertion into the body and having a surface providing said interface .

Suitably, the unclad portion is at a distal end of said fibre and has a reflective cap, e . g . of gold or silver . Suitably, the optical fibre is a single mode silver halide (AgBr_xCl_Lx) fibre .

It can be envisaged that this device may also be used for non-invasive measurement of analyte concentration, for example, by placing the unclad portion of the fibre against the skin of the subj ect . Lastly, in a sixth aspect of the invention, there is provided optical apparatus- having a source of optical radiation and having first and second optical paths therefrom to respective detectors , wherein said source is a resistance heating element and said paths originate at said source and diverge therefrom, preferably in opposite directions .

The invention will be further described and illustrated with reference to the accompanying drawings , in which :

Figure 1 shows a schematic side view of the principal components of a non-invasive analyte concentration measurement device according to the present invention .

Figure 2 is a plan view from above corresponding to Figure 1.

Figure 3 shows the angle of incidence dependence of the transmission of silica collimator tubes for two wavelengths . Figure 4 shows the penetration depth of the evanescent wave from a ZnSe crystal into tissue having refractive index 1.273 , for a selection of wavelengths in the range 5 - 10 |im .

Figure 5 shows the absorbance spectrum of glucose over a wavelength range of 5 to 12 μm.

Figure 6 shows the IR spectrum of water measured at 25 ⁰C over the wavelength range from 2.5 μm to 40 μm .

Figure 7 shows a schematic diagram of the optics of an invasive analyte concentration measurement device according to the present invention .

Figure 8 shows a schematic diagram of the optical waveguide fibre of the invasive analyte concentration measurement device of Figure 7.

Figure 9 shows a comparison of the glucose concentrations obtained for a non-diabetic subj ect using a standard finger- prick determination method and the apparatus of the present ■ invention; and

Figure 10 shows similar results obtained from a diabetic subject .

The principal components of a non-invasive analyte concentration measurement device according to an embodiment of the present invention which reflects all of the different aspects of the invention are shown schematically in Figure 1 and Figure 2. Figure 1 shows the device in side view and Figure 2 shows the device in plan view, both at an approximately 1 : 1 scale .

A source of electromagnetic radiation 1 is provided which comprises a thin strip of Nichrome™ foil . Passage of electrical current through this strip causes broadband electromagnetic radiation to be produced from both sides of the foil . Thus , the foil is capable of emitting radiation in both a forward and a backward direction. The source is connected to electrical circuitry (not shown) comprising a capacitor in a power supply circuit powered by a 12 V battery by which a pulse of radiation may be produced each time the capacitor discharges through the strip of foil . Use of a battery as power source improves the portability of the device, which is an important factor for a device intended for regular use by a patient . The ability of the emitter to emit radiation in two directions avoids the need to employ a beam splitter in order to monitor the incident flux - the emitter is instead directly monitored . This reduces the cost of the device significantly, again an important consideration for a device for home use by a patient . The backward-emitted radiation from the source enters a collimator 2 , which directs the radiation towards a dual-channel detector system 3. The collimator 2 is a hollow waveguide e . g . one comprised of hollow silica ( SiO₂ ) , which may be coated internally with a silver mirror finish and silver halide . The hollow silica may be in the form of circular cross section tubes . Typical dimensions of such tubes are : internal diameter 3.2 mm, external diameter 4.8 mm, length 25 mm. Such tubes may be obtained from several commercial suppliers , such as Quartz Scientific , Inc . Alternatively, another shape of hollow silica collimator may be used, such as a cuboid of rectangular cross section . One such cuboid collimator may replace the two circular tubes illustrated. Silica glass has an unusual refractive index of less than 1 in the wavelength range 7 - 10 μm; the difference in the _ resultant transmission for light of 5 μm and below and light of 9 μm is shown in Figure 3. It can be seen from Figure 3 that the transmission through the silica tube is significantly greater over a range of angles of incidence of light at 9 μm compared to 5 μm. The efficiency with which the light reaches the detector is therefore improved compared with conventional collimators , and so a less powerful source may be used, allowing the power requirement to be reduced to that which may be supplied by a battery. Alternative materials having a refractive index of less than one in the IR region around 10 μm may also be used for the collimator tubes , such as GeO₂. The expense of using silica tubes is considerably less , than that of a conventional lens arrangement to focus the light on to the detector . The size of the device may also be significantly reduced as the focal length of suitable lenses is relatively long compared with the necessary length of the collimator tubes ( ZnSe lenses have a focal length of at least 10 cm) .

The detector system 3 to which the backward-emitted radiation is directed by the silica collimator is a dual channel pyroelectric detector , comprising two piezoelectric crystal detectors , each having a filter to restrict the radiation incident on the crystals to the desired wavelength range for the measurement and reference channels , respectively . Typically, the filters used are those supplied with the detectors by the commercial supplier . For example , a Spectrogen filter (BP-2200- 100-B) may be used to select wavelengths around 2.2 μm. The pyroelectric detector system does not require cooling and is suitable for use in a low-cost device . The detector system 3 monitors the intensity of the backward emission of- the radiation source , as a reference .

Forward emission from the source 1 passes into one end of an ATR crystal 4. This crystal is preferably a zinc selenide

( ZnSe) prism, as zinc selenide has a high refractive index ( 2.4 at 10.6 μm) and is inert towards the skin . As an alternative, silver halide may be used, although this is not preferred due to its light sensitivity. If desired, the crystal may be capped with CVD diamond, which has a matching refractive index, in order to protect the external surfaces of the crystal from abrasive damage . The thickness of the crystal is around 2 mm so as to provide at least approximately 25 reflections within the interface area at which measurement is conducted. Typically, the plate may be 52 mm long and 20 mm wide . Two long sides of the crystal should be substantially parallel , and one of those sides should be ^'that against which the skin surface is placed. This last surface may incorporate marked areas 7 onto which the skin surface ( s ) should be placed, for example, fingertips . The areas surrounding the marked areas optionally are masked so as not to allow the passage of evanescent waves . A suitable masking material may be non-toxic glass-adhering enamel . More preferably, a film of gold may be deposited on the crystal to form the masked areas . The temperature of this surface may be monitored by a temperature measuring device such as thermistor 8 , and/or a pressure sensor (not shown) . The electromagnetic radiation passes along the crystal by means of total internal reflection, and exits the crystal at the far end from the source . The radiation incident on the areas of the crystal surface in contact with the skin surface produces an evanescent wave which penetrates a small distance into the skin, before being reflected; the reflected beam carries an absorption spectrum of the tissue penetrated by the evanescent wave .

The radiation exiting, the crystal passes into collimator tubes 5 which may take the same form as the collimator 2 and which serve to direct the radiation to a detector system 6 , which is as described for the detector system 3 above . The angle φ at which the detector system 6 views the interface surface of the crystal 4 may be altered in order to optimise capture of the radiation from the source . Physically, this is accomplished by mounting the length of this unit (tubes plus detector system) on a suitable pivoting carriage (not shown) . Changing the angle of view of the detector system 4 has the effect of selecting rays for detection that impinge on the interface at different angles of incidence from the collimator tubes 2. This angle of incidence determines the depth of penetration of the evanescent wave into the skin surface, as shown in Figure 4. The optimum penetration depth is that at which there is maximum detected absorbance of the forward- emitted radiation and hence maximum signal to noise ratio; the angle of incidence required to attain this maximum will vary from person to person according to their skin type . The optimum angle may therefore be determined by trial and error for a specific person and may then be fixed for future use by that person . If the apparatus is for use by several persons , the individual settings for each person may be held in a suitable database . The movement of the detector system 4 to alter the viewing angle may be motorised.

The apparatus includes a computation unit 8 which has input means such as a keyboard for entering calibration data . The wavelength ranges for the measurement and reference channels detected by the detectors are chosen by inspection of the absorption spectrum of glucose . Mid-IR light in the range 6.5 to 14 μm is found to be absorbed by glucose , with the intensity of the absorption having a strong dependency on the concentration of glucose present ( see Figure 5 ) . The spectrum of water is featureless in this region, and sweat on the skin does not to affect the measurement obtained ( see Figure 6 ) . It can be seen from Figure 5 that there is significant absorption by glucose at the wavelength band of 8.25 - 8.75 μm (wavenumbers 1142 - 1212 ) suggested by Berman . The measurement wavelength range preferred by the inventors is 6.6 to 14 μm; alternatively, the range 8.8 to 10.5 Um may be used. The reference wavelength range may be any region in which the intensity of the absorption spectrum of glucose is substantially independent of the glucose concentration. Preferably, the reference wavelength range is a range of shorter wavelengths than the measurement wavelength range, as shorter wavelengths penetrate less far into the skin surface ( see Figure 4 ) , and therefore respond more sensitively to surface perturbations and contarru nation of the skin whilst also having less exposure to glucose . This may be used to correct for such contamination in the measurement result . As the reference wavelength λ_ref is in the denominator of the ratio expression used to calculate the glucose concentration (vide infra) , a smaller λ_reE is preferred to increase the signal-to- noise ratio for the measurement . The preferred reference wavelength ranges are 5.0 to 5.6 μm and 1.8 to 2.6 Um.

Calculation of the analyte concentration from the measured absorptions of the reference and measurement wavelengths from the forward and backward emitted radiation is performed using a ratio method, as follows :

A finger of refractive index n₂ is placed on a ZnSe crystal of refractive index nι . Light of wavelength say 5 to 15 μm is incident on the crystal . The light is transmitted through the crystal by internal reflection . Some of the light penetrates the finger where it is attenuated by glucose in the tissue before being reflected into the crystal .

The depth to which the light penetrates the finger, δ, is small , and is given by the expression :

where ^ = angle of incidence or angle of reflection

The effective path length of the light , 1 , in the finger is given by: I = pλ where p is a constant for fixed θ, τiχ and n₂ such that

The glucose concentration in the tissue of the finger can be derived as follows :

The loss at each reflection in the tissue is T - T _e-P^te where I₀ is the incident light intensity and ε is the absorption coefficient The intensity of the light , J, after a single reflection from the finger is / = I_oe-^pλε and after N reflections , the reflected light intensity I_N will be

N depends on the contact area of the finger with respect to the crystal , and will vary with variation of the finger pressure on the crystal . To reduce this dependence, the finger pressure is kept constant . We choose two wavelengths : λ_ref, where no change in ε with glucose concentration occurs , and λ_me_as where a change in ε with glucose concentration occurs . Thus ,

^«£ ^NPKet ^εr_ef

and

metis

Therefore , the ratio of the two intensity ratio logarithms at the reference and measurement wavelengths is

Ratio = man meai

Let ε_ref = s_ref (absorption coefficient for skin etc ) ; and

^£mea\ ~ ^Smeas ^W-¹

where g is the concentration of glucose and G is the additional absorption due to glucose .

Thus , the ratio of the intensity ratio logarithms becomes

which may be written in the form Ratio = D + Eg. This expression does not depend on N, and thus is independent of the contact area . The ratio is linear in g. Thus , D and E must be known to determine the glucose concentration, and a convenient way to do this is to determine D and E at low and high glucose concentration . ε_ref and E_meas can be determined, which yields D and E. The system is thus calibrated to measure glucose concentrations non-invasively.

Alternatively, let : S_ref = S₁ + XW₁ and e_ms3S = S₂ + xW₂ + gG, where : S₁ is the absorption coefficient of skin etc at the reference wavelength;

S₂ is the absorption coefficient of skin etc at the measurement wavelength; W₁ is the absorption coefficient of water at the reference wavelength;

W₂ is the absorption coefficient of water at the measurement wavelength;

G is the absorption coefficient due to glucose ; x is the concentration of water in the region of the body part probed by the radiation,- g is the concentration of glucose in mmol . dirf³. Then :

d = = {s₂ +xW₂)kl₂ +gGkl₂ ,

where I_x is the effective path length at the reference wavelength, Z₂ is the effective path length at the measurement wavelength, and k is a constant .

Manipulation of these expressions leads to the result that :

where F(c) is a function of c, which is best expressed as

7 r— , giving the expression g = (2), but for values

of { d-c) of greater than about 0.45 is proportional to , so d -c that one may use the expression

where A ' is a constant for given conditions .

We find that A '/A is approximately 13 /4.

For lower values of (d-c) , the proportionality between

no longer holds , but a polynomial expression

F(d-c) , which rises more rapidly as (d-c) falls than does the

expression

This expression ( 2 ) does not depend on JV, and thus is independent of contact area . A is a constant for a given angle , pressure and temperature; A must be known to determine the glucose concentration . This may be achieved by determining a glucose concentration g, for example from a blood sample in a conventional manner , and inputting this into the apparatus in order that A may be determined using expression ( 2 ) . Experimentally, it is found that A varies little from subj ect to subj ect as a consequence of differences in refractive index n. An independent determination of n could lead to a fully self- calibrated device , as :

, W_n(a+x) W₇(a+x)l_{ A = —= = , and, as stated above, / is

G Gl₂ related to the refractive index n through the expression / = pλ

The pressure applied by the skin surface to the surface of the ATR crystal should remain constant above a pre-determined minimum pressure in order to keep the optical properties of the skin consistent . This also keeps N constant , which is desirable even though the above expression is independent of N. This may be achieved in two principal ways . A pressure sensor may be incorporated into the surface of the ATR crystal in order to determine the pressure applied by the skin surface . A feedback system may also be incorporated so that if at least the desired minimum pressure is not maintained a signal may alert the user to apply more pressure . As an alternative, or in combination with the above pressure sensor , a pressure-maintaining member may be comprised in the device , which may apply pressure to the opposite side of the body part from that in contact with the ATR crystal in order that at least the desired minimum pressure is maintained.

It is observed that the constant A in expression ( 2 ) above shows a dependence on ambient temperature . The sensitivity of A to temperature may be greater in the diabetic patient as the fingers of diabetic patients are often poorly perfused. The change in the value of A with temperature may be characterised in order to obtain a value of A corrected for the temperature at which the measurement is taken for use in equation ( 1 ) . Preferably, the change in the value of A with temperature is characterised using the expression:

A(T)=A₀[I-_P(T-T₀)], where : A (T) is the value of A at temperature T; A₀ is the value of A at a reference temperature T₀; p is a constant for a given set of conditions of use and subj ect and is found to have a value of approximately 0.2. The value of p may be larger for diabetic patients due to differences in peripheral perfusion .

Further, it is preferred to monitor the effect of changing water concentration in the stratum corneum to determine an equilibrium value for the analyte concentration reading . When measuring glucose, it is seen that the measured glucose concentration changes over an initial period of some minutes .

This may be due to changes in the water concentration consequent upon the skin being placed on the ATR plate . The change in measured concentration may be modelled using a simple exponential law, and a regression comparing the least squares exponential fit with the average of a fixed number of scans may be carried out until a satisfactory convergence has been reached . Preferably, the Levenberg-Marquardt algorithm is used to perform the regression . The values of d and c may then be recorded. The values of a and b may also be recorded from time to time, preferably once a day. This determination may be carried out by placing water or other aqueous calibration medium directly on to the ATR plate in the measuring areas instead of the fingertips , and carrying out a concentration determination in the normal manner . However, for a given apparatus the ratio b/a is well approximated by a constant value . For instance , for the apparatus just described it is approximately 4.17. Also, if two different calibration measurements of blood glucose are taken, they can be used to determine both A and b/a .

For the measurements obtained using these regressions to be reliably accurate, it is necessary that the correlation of the measured data with the function to which it is approximated is monitored, and if the correlation becomes invalid, for example due to movement of the subject , the measurement process must be repeated . Preferably, therefore , the correlations of both regressions are monitored. More preferably, if the correlations drop below pre-chosen levels at any time , the software prompts for appropriate action, for example : cancel , ignore, continue, etc .

Although principally intended for use as described above with a body portion effectively constituting a sample for measurement , the illustrated apparatus can be used in a modified method in which a sample such as a liquid containing an analyte is directly placed on the surface of the ATR prism 4.

The principal components of an invasive analyte concentration measurement device are shown in Figure 7 and Figure 8.

Referring to Figure 7 , an optical arrangement for an invasive glucose monitor according to the present invention is shown . A two-sided emitter.21 as described above for the non- invasive device is used as the source of broadband electromagnetic radiation . Preferably, the radiation is emitted as pulses generated as in the previous case . The backward- emitted radiation is directed towards a pyroelectric detector system 26 by silica collimator tubes 22 , as described previously, to provide null absorption spectra of the chosen reference and measurement wavelength ranges . The forward- emitted radiation passes through a zinc selenide convex lens 30 , through a KBr beam splitter 32 and then through a second convex lens 34 , which focuses the radiation into a proximal , uncapped end of an optical fibre 24. The radiation is then transmitted along the optical fibre by total internal reflection, with attenuation caused by the absorptions of components of the surrounding medium where the fibre is unclad ( see below) . At the capped end, the radiation is reflected and passes back along the fibre to exit the open end, whence it passes through the lens 34 , and is split by the beam splitter 32. Part of the radiation is focused onto a detector system 37 by means of a third convex lens 36 , and the absorption spectra of the chosen measurement and reference wavelength ranges are again recorded. The detector 37 is a two-channel pyroelectric detector incorporating two filters , one for each chosen wavelength range , as described for the non-invasive device .

Referring to Figure 8 , the fibre 24 is shown in more detail . The preferred material for the optical fibre is single mode silver halide (AgBr_xCli__x) ; this material is non-toxic, biocompatible, has a refractive index of 2.2 at 10.6 μm, and will transmit wavelengths in the range 2 to 30 μm. Such a single mode fibre is particularly preferred as no change in mode occurs when the fibre is bent . It is preferred that a proximal portion 38 of the fibre 24 is clad with a material that does not allow the evanescent wave to interact with the tissue surrounding the fibre , and that a distal part 40 of the fibre 24 is unclad such that the evanescent wave may interact with the surrounding medium. A suitable material for cladding the fibre is polycarbonate . The distal end of the fibre is capped with a reflective material , preferably a bio-compatible, non-toxic material , such as gold, silver, or aluminium .

As described previously, the reference wavelength range may be selected to be in any region of the absorption spectrum ^' of the analyte of interest where the intensity of the absorption is substantially independent of the analyte concentration . The measurement wavelength range may be selected to be in any region of the absorption spectrum of the analyte wherein the intensity of the absorption is significantly dependent on the analyte concentration . The wavelength range in which the measurement and reference wavelengths preferably fall is 5 to 14 μm. Over this range of wavelengths , the spectrum of water is relatively featureless . The analyte concentration may be measured using the ratio method described above , or the method using expression (2 ) described above .

The optical fibre may be placed in a body region such that the unclad area of the fibre is in contact with the tissue in which the analyte of interest is found . Alternatively, the device may be placed in contact with a sample solution containing the analyte of interest . The area of unclad fibre is limited in order that the number of reflections N occurring in contact with the sample is kept constant during the measurements . It is also possible to place the unclad portion of the optical fibre against a skin surface, for example, held between two fingers of the subject, and thereby to use this device as a non-invasive glucose sensor . If the device is used in this manner, it is preferable also to perform the appropriate temperature compensation described above .

Example 1 - Non-Invasive Glucose Concentration Measurement

This example refers to the use of the apparatus of Figure 1. Initial Calibration After overnight fasting , the subj ect removes the outer layer of the stratum corneum of the index fingers by applying and peeling off strips of adhesive tape five times . The subj ect then washes the index fingers in water , dries them, then cleans them with isopropanol and allows them to dry in air . The plate of the ATR crystal on to which the fingers will be placed is also cleaned with isopropanol and allowed to air dry. The instrument is allowed to complete its self-calibration . This self-calibration comprises the device recording the temperature and the ratio of the direct (I₀) to the indirect ( I_N) intensities in the absence of the subj ect ' s fingers for each of 10 angles .

The subject then places his index fingertips on the marked regions of the ATR plate . The device will record the absorption ratio and the value of the null channel having driven the motorized detector stage to match the angle of the detector 6 with the ATR crystal 4 in order to optimise performance . The device will then prompt the subject to remove his fingers from the plate .

The subj ect then takes a blood sample from a finger other than the index finger and measures his blood glucose concentration in the usual way, noting the reading . He then raises his blood sugar level by eating or drinking, and after an appropriate time interval takes a second blood sample and glucose concentration reading in the conventional manner .

The subj ect then places his index fingertips on the marked regions of the ATR crystal and repeats the measurement process above . The two values of the blood sugar levels from the conventional determinations may then be keyed into the ATR device, and the device will use these data to calculate the calibration constants as describe above . Regular Measurement Routine The subj ect should clean his fingers as described above, and place his index fingers on the marked parts of the plate of the ATR crystal until prompted to remove them. The device will use the previously calculated calibration constants to produce a reading for the tissue glucose concentration .

Usage Example 2 - Non-Invasive Glucose Concentration Measurement

This example refers to the use of the apparatus of Figure 1. Initial calibration The subj ect removes the outer layer of the stratum corneum of the index fingers by applying and peeling off strips of adhesive tape five times . The subj ect then washed the index fingers in water, dries them, then cleans them with isopropanol and allows them to dry in air . The plate of the ATR crystal on to which the fingers will be placed is also cleaned with isopropanol and allowed to air dry. The instrument is allowed to complete its self-calibration, as above.

The subj ect takes a blood sample from a finger other than the index finger and measures his blood glucose concentration in the usual way. The reading obtained is keyed into the ATR device in order to calculate a value of A .

The subject then places his index fingertips on each of the marked regions of the ATR plate and lowers the pressure maintaining member over his fingers . The device will record the values of d and c , having driven the motorised detector stage to match the angle of the detector 6 with the ATR crystal 4 in order to optimise performance . The device will then prompt the subj ect to remove his fingers from the plate .

From time to time, water should be placed on the ATR plate in the marked areas , and the device allowed to measure a and b . Regular Measurement Routine

The subj ect should clean his fingers as described above, and place his index fingers on the marked parts of the plate of the ATR crystal until prompted to remove them . The device will use the previously determined value of A to produce a value for the tissue glucose concentration .

A comparison between glucose concentration values obtained using a conventional method of determining blood glucose and use of the method described in this Example is shown in Figure 9.

Usage Example 3 - Invasive Glucose Concentration Measurement

This example refers to the use of the apparatus of Figure 7. A qualified medical practitioner inserts a catheter just below the skin of the patient . The fibre , having been cleaned with isopropanol and sterilised, may then be fed through the catheter to a calibrated distance . The measurement routine may then be carried out by the instrument as given in either of the non-invasive examples .

Claims

1. Apparatus for measuring analyte concentration in a living body, comprising a source of optical radiation, an optical waveguide having a surface for forming an interface with a body region of said living body, and at least a first detector system, comprising at least one detector , for quantitative detection of at least a measurement wavelength band of said radiation for which an absorption characteristic of said analyte is dependent on the analyte concentration adjacent said optical waveguide surface in contact with said body region to produce an optical absorbance measure d and for quantitative detection of at least a reference wavelength band of said radiation to produce an optical absorbance measure c, wherein said source of optical radiation is positioned to supply said radiation to said optical waveguide to interact as an evanescent wave with the body region containing the analyte via the interface between the body region and the optical waveguide, and the or each detector of the first detector system is positioned to receive said radiation from said optical waveguide after said interaction, and wherein the reference wavelength band is selected to be in a wavelength range for which an absorption characteristic of said analyte is substantially less dependent on the analyte concentration, and computation means for receiving and processing an electronic output provided by at least said first detector system representing said quantitative detection of said measurement and reference wavelengths of said radiation to produce an output indicative of the concentration of said analyte , wherein said output is calculated as being directly proportional to an arithmetical combination of said quantitative measures d and c and a factor b/a equivalent to

wherein g is the analyte concentration, A is constant for a given set of conditions of use , F (c) is a function of c, and b/a is constant for a given set of conditions of use .

2. Apparatus as claimed in claim 1 , wherein said function

F (c) is so that g is calculated as an

arithmetical combination of said quantitative measures d and c and a factor b/a equivalent to

3. Apparatus as claimed in claim 1 or claim 2 , wherein said computation means is configured for receiving and processing an electronic output provided by said detector system representing optical absorbance values b and a for said measurement and reference wavelength bands respectively, and corresponding to measures d and c respectively, but which are determined with an aqueous reference material rather than said body region in contact with said optical waveguide surface .

4. Apparatus as claimed in claim 1 or claim 2 , wherein said computation means is configured for receiving utilising a calculated or separately determined value of b/a .

5. Apparatus as claimed in claim 4 , wherein said calculated or separately determined value of h/a, represents an estimate of the ratio of optical absorbance values b and a which would be produced for said measurement and reference wavelength bands respectively if the apparatus were to be operated with an aqueous reference material rather than said body region in contact with said optical waveguide surface .

6. Apparatus as claimed in any one of claims 3 to 5 , wherein said aqueous reference material is water .

7. Apparatus for measuring analyte concentration in a living body, comprising a source of at least mid-IR radiation, an optical waveguide having a surface for forming an interface with a body region of said living body, and at least a first detector system, comprising at least one detector, for quantitative detection of at least a measurement wavelength band of said radiation for which an absorption characteristic of said analyte is dependent on the analyte concentration adj acent said interface and for quantitative detection of at least a reference wavelength band of said radiation, wherein said source of mid-IR radiation is positioned to supply said radiation to said optical waveguide to interact as an evanescent wave with the body region containing the analyte via the interface between the body region and the optical waveguide, and the or each detector of the first detector system is positioned to receive said radiation from said optical waveguide after said interaction, and wherein the reference wavelength band is selected to be in a wavelength range for which an absorption characteristic of said analyte is substantially independent of the analyte concentration.

8. Apparatus for measuring an analyte concentration in a living body, comprising a source of at least mid-IR radiation, an optical waveguide having a surface for forming an interface with a body region of said living body, a temperature measuring device measuring the temperature of the body region at the interface , and at least a first detector system, comprising at least one detector, for quantitative detection of at least a measurement wavelength band of said radiation^' for which an absorption characteristic of analyte is dependent on the analyte concentration adj acent said interface, and for quantitative detection of at least a reference wavelength band of said radiation, wherein said source of mid-IR radiation is positioned to supply said radiation to said optical waveguide to interact as an evanescent wave with the body region containing the analyte via the interface between the- body region and the optical waveguide, and the or each detector of the first detector system is positioned to receive said radiation from said optical waveguide after said interaction, the detector system and the temperature measuring device providing electrical inputs to a computer adapted to compute an analyte concentration compensated for the temperature of the body region.

9. Apparatus as claimed in claim 8 , wherein said compensation comprises adjustment • of the value of the constant A based on the equation

A(T)=A₀[I-_P(T-T₀)]

where T is the temperature in degrees Celsius , p is a constant , and A₀ indicates the value of A at any reference temperature T₀.

10. Apparatus as claimed in any preceding claim, wherein the optical waveguide has a light exit face from which said radiation exits after said interaction with the analyte and wherein said optical waveguide and said first detector system are mounted for relative movement such that an angle at which said light exit face is viewed by the detector system is variable by relative movement of the detector system and the optical waveguide .

11. Apparatus for measuring analyte concentration in a sample , comprising a source of electromagnetic radiation emitting said radiation in at least a first direction, an optical waveguide having a surface for providing an interface with a sample" placed in contact therewith and positioned to receive said electromagnetic radiation emitted in said first diz-ection to interact: as an evanescent wave with a said sample via said interface between the sample and the optical waveguide surface at a plurality of different angles of incidence with said interface, at least a first detector system for detecting said electromagnetic radiation emitted in said first direction after passage through said waveguide , wherein the said first detector system is mounted for relative movement with respect to a light exit face of the optical waveguide from which the electromagnetic radiation exits the waveguide after interaction with the sample such that an angle at which the light exit face is viewed by the detector system is variable by relative movement of the detector system and the optical waveguide to thereby select for detection the radiation which has a desired angle of incidence to said interface .

12. Apparatus as claimed in any preceding claim, comprising a second detector system comprising at least one detector positioned to receive the radiation output from said source without passage through said optical waveguide or without interaction at said interface .

13. Apparatus for measuring analyte concentration in a sample , comprising a source of electromagnetic radiation emitting said radiation in at least a first direction and a second direction, an optical waveguide having a surface for providing an interface with a sample placed in contact therewith and positioned to receive said electromagnetic radiation emitted in said first direction to interact as an evanescent wave with a said sample via said interface between the sample and the optical waveguide surface , at least a first detector system comprising at least one detector for detecting, said electromagnetic radiation • emitted in said first direction after passage through said'^" waveguide , a second de.tector system comprising at least one detector for detecting said electromagnetic radiation emitted in said second direction, each said first detector system providing an electronic output , and a computational circuitry for receiving and processing said electronic outputs to derive therefrom the concentration of the analyte .

14. Apparatus as claimed in claim 13 , wherein said first and second detector systems each quantitatively detects at least a measurement wavelength band of said radiation for which an absorption characteristic of said analyte is dependent on the analyte concentration adjacent the interface and quantitatively detects at least a reference wavelength band of said radiation .

15. Apparatus as claimed in claim 13 or claim 14 , wherein the optical waveguide has a light exit face from which said radiation exits after said interaction with the analyte and wherein said optical waveguide and said first detector system are mounted for relative movement such that an angle at which said light exit face is viewed by the detector system is variable by relative movement of the detector system and the optical waveguide .

16. Apparatus as claimed in any preceding claim, wherein the optical waveguide comprises an optical fibre having a proximal clad portion and a distal unclad portion having a surface providing said interface .

17. Apparatus as claimed in claim 16 , wherein the optical fibre is a single mode silver halide (AgBr_xCli__x) fibre .

18.: Apparatus for measuring analyte concentration in a living body, comprising a source of optical radiation, an optical waveguide having a surface for forming an interface with a skin region of said living body, and at least a first detector system, comprising at least one detector, for quantitative detection of at least a measurement wavelength band of said radiation for which an absorption characteristic of said analyte is dependent on the analyte concentration adjacent said optical waveguide surface in contact with said skin region to produce an optical absorbance measure d and for quantitative detection of at least a reference wavelength band of said radiation to produce an optical absorbance measure c , wherein said source of optical radiation is positioned to supply said radiation to said optical waveguide to interact as an evanescent wave with the body region containing the analyte via the interface between the skin region and the optical waveguide , and the or each detector of the first detector system is positioned to receive said radiation from said optical waveguide after said interaction, and wherein the reference wavelength band is selected to be in a wavelength range for which an absorption characteristic of said analyte is substantially less dependent on the analyte concentration, and computation means for receiving and processing an electronic output provided by said detector system representing said quantitative detection of said measurement and reference wavelengths of said radiation to produce an output indicative of the concentration of said analyte, wherein said electronic output is dependent on time elapsed from the placing of the skin region on said surface and said computation means operates to predict a maximum value for the measured analyte .concentration from ■ values of said output obtained prior to said output reaching a stable level . ^• - ^{■ ■}'.. ' ^• .

19. Apparatus as claimed in claim 17 , wherein predicted maximum values for the analyte concentration are repeatedly calculated and the rate of change in said predicted maximum values is determined, and wherein a reading of analyte concentration is output when said rate of change falls below a preset level .

20. Apparatus as claimed in claim 18 or claim 19 , wherein the optical waveguide has a light exit face from which said radiation exits after said interaction with the analyte and wherein said optical waveguide and said first detector system are mounted for relative movement such that an angle at which said light exit face is viewed by the detector system is variable by relative movement of the detector system and the optical waveguide .

21. Apparatus as claimed in any one of claims 1 to 17 , for measuring an analyte concentration in the skin of a living body .

22. Apparatus as claimed in any one of claims 18 to 21 , wherein a pressure-maintaining member is provided to maintain adequate pressure of the body region against the said interface , and/or wherein the pressure applied to the said interface is measured by a pressure sensor .

23. Apparatus as claimed in any one of claims 7 to 22 , comprising computation means for receiving and processing an electronic output provided by at least said first detector system representing said quantitative detection . of said measurement and reference wavelengths ^■ of said radiation to produce an output indicative of the concentration of said analyte, wherein said output is _' •^■ . calculated as being directly proportional to an^' arithmetical combination of said quantitative measures d and c and a factor b/a equivalent to

wherein g is the analyte concentration, A is constant for a given set of conditions of use, F(c) is a function of c, and b/a is constant for a given set of conditions of use .

24. Apparatus as claimed in claim 23 , wherein said function

F(c) is SO that g is calculated as an

25. Apparatus as claimed in claim 23 or claim 24 , wherein said computation means is configured for receiving and processing an electronic output provided by said detector system representing optical absorbance values b and a .for said measurement and reference wavelength bands respectively, and corresponding to measures d and c respectively, but which are determined with an aqueous . reference material rather than said body region in contact with said optical waveguide surface .

26. Apparatus as claimed in claim 23 or claim 24 , wherein said computation means is configured for receiving utilising a calculated or separately determined value of b/a .

27. Apparatus as claimed in claim 26 , wherein said calculated or separately determined value of b/a, represents an estimate of the ratio of optical absorbance values b and a which would be produced for said measurement and reference wavelength bands respectively if the apparatus were to be operated with an aqueous reference material rather than said body region in contact with said optical waveguide surface .

28. Apparatus as claimed in any one of claims 25 to 27 , wherein said aqueous reference material is water .

29. Apparatus as claimed in any preceding claim, wherein the optical waveguide conducts said radiation by multiple total internal reflections at said surface providing said interface such that light is reflected at said surface at least 25 times in passing through said waveguide .

30. Apparatus as claimed in claim 1-15 or 17-29 , wherein said optical waveguide is in the form of a plate or is in the form of an optical fibre .

31. Apparatus as claimed in any one of claims 1 to 10 or 14 to 30 , wherein the measurement wavelength band is in the mid-IR region and wherein the reference wavelength band is in the mid-IR, near-IR, or visible region.

32. Apparatus as claimed in claim 31 , wherein the measurement wavelength band is in the range 6.6 to 14 μm.

33. Apparatus as claimed in claim 31 or claim 32 , wherein the reference wavelength is in the range 5.0 to 5.6 μm.

34. Apparatus as claimed in claim 31 or claim 32 , wherein the reference wavelength is in the range 1.8 to 2.6 μm.

35. Apparatus as claimed in any one of claims 1 to 10 or 14 to 34 , wherein the absorption coefficient of the analyte at the reference wavelength band is no greater than 1% when measured on a solution of the analyte in a suitable solvent .

36. Apparatus as claimed in any one of claims 1 to 10 or 14 to 35 , wherein the ratio of the absorption coefficient of the analyte at the measurement wavelength band to the absorption coefficient of the analyte at the reference

5- wavelength band is greater than 99% when measured on a solution of the analyte in a suitable solvent .

37. Apparatus as claimed in any preceding claim, wherein the or each said detector system comprises a measurement

10 optical filter selectively passing said measurement wavelength band and a reference optical filter selectively passing said reference wavelength band .

38. Apparatus as claimed in claim 37 , wherein the or each 15 detector system comprises respective measurement and reference detectors positioned to receive the filtered output of said measurement and reference optical filters respectively.

20.

39. Apparatus as claimed in any preceding claim, wherein the radiation is directed from the source towards the optical waveguide and/or from the optical waveguide to the or a detector system, using collimator tubes having a refractive index of less than one in the wavelength region

25 around 10 μm.

40. Optical apparatus having a source of optical radiation and having first and second optical paths therefrom to respective detectors , wherein said source is a resistance

30 heating element and said paths originate at said source and diverge therefrom.

41. Apparatus as claimed in claim 40 , wherein said optical paths are in opposite directions .