The need and demand for an accurate, non-invasive method for determining blood glucose level in patients is well documented. Barnes et al. (U.S. Pat. No. 5,379,764) disclose the necessity for diabetics to frequently monitor glucose levels in their blood. It is further recognized that the more frequent the analysis, the less likely there will large swings in glucose levels. These large swings are associated with the symptoms and complications of the disease, whose long term effects can include heart disease, arteriosclerosis, blindness, stroke, hypertension, kidney failure, and premature death. As described below, several systems have been proposed for the non-invasive measurement of glucose in blood. However, despite these efforts a lancet cut into the finger is still necessary for all presently commercially available forms of home glucose monitoring. This is believed so compromising to the diabetic patient that the most effective use of any form of diabetic management is rarely achieved.
The various proposed non-invasive methods for determining blood glucose level, discussed individually below, generally utilize quantitative infrared spectroscopy as a theoretical basis for analysis. Infrared spectroscopy measures the electromagnetic radiation (0.7-25 .mu.m) a substance absorbs at various wavelengths. Molecules do not maintain fixed positions with respect to each other, but vibrate back and forth about an average distance. Absorption of light at the appropriate energy causes the molecules to become excited to a higher vibration level. The excitation of the molecules to an excited state occurs only at certain discrete energy levels, which are characteristic for that particular molecule. The most primary vibrational states occur in the mid-infrared frequency region (i.e., 2.5-25 .mu.m). However, non-invasive analyte determination in blood in this region is problematic, if not impossible, due to the absorption of the light by water. The problem is overcome through the use of shorter wavelengths of light which are not as attenuated by water. Overtones of the primary vibrational states exist at shorter wavelengths and enable quantitative determinations at these wavelengths.
It is known that glucose absorbs at multiple frequencies in both the mid-and near-infrared range. There are, however, other infrared active analytes in the blood which also absorb at similar frequencies. Due to the overlapping nature of these absorption bands, no single or specific frequency can be used for reliable non-invasive glucose measurement. Analysis of spectral data for glucose measurement thus requires evaluation of many spectral intensities over a wide spectral range to achieve the sensitivity, precision, accuracy, and reliability necessary for quantitative determination. In addition to overlapping absorption bands, measurement of glucose is further complicated by the fact that glucose is a minor component by weight in blood, and that the resulting spectral data may exhibit a non-linear response due to both the properties of the substance being examined and/or inherent non-linearities in optical instrumentation.
A further common element to non-invasive glucose measuring techniques is the necessity for an optical interface between the body portion at the point of measurement and the sensor element of the analytical instrument. Generally, the sensor element must include an input element or means for irradiating the sample point with infrared energy. The sensor element must further include an output element or means for measuring transmitted or reflected energy at various wave lengths resulting from irradiation through the input element.
Robinson et al. (U.S. Pat. No. 4,975,581) disclose a method and apparatus for measuring a characteristic of unknown value in a biological sample using infrared spectroscopy in conjunction with a multivariate model that is empirically derived from a set of spectra of biological samples of known characteristic values. The above-mentioned characteristic is generally the concentration of an analyte, such as glucose, but also may be any chemical or physical property of the sample. The method of Robinson et al. involves a two-step process that includes both calibration and prediction steps. In the calibration step, the infrared light is coupled to calibration samples of known characteristic values so that there is differential attenuation of at least several wavelengths of the infrared radiation as a function of the various components and analytes comprising the sample with known characteristic value. The infrared light is coupled to the sample by passing the light through the sample or by reflecting the light from the sample. Absorption of the infrared light by the sample causes intensity variations of the light that are a function of the wavelength of the light. The resulting intensity variations at the at least several wavelengths are measured for the set of calibration samples of known characteristic values. Original or transformed intensity variations are then empirically related to the known characteristic of the calibration samples using a multivariate algorithm to obtain a multivariate calibration model. In the prediction step, the infrared light is coupled to a sample of unknown characteristic value, and the calibration model is applied to the original or transformed intensity variations of the appropriate wavelengths of light measured from this unknown sample. The result of the prediction step is the estimated value of the characteristic of the unknown sample. The disclosure of Robinson et al. is incorporated herein by reference.
Several of the embodiments disclosed by Robinson et al. are non-invasive and incorporate an optical interface having a sensor element. As depicted in FIGS. 5 and 6 of Robinson et al., the optical interface includes first, an input element and second, an output element. The input element is an infrared light source or near infrared light source. The input element interface with the sample or body portion containing blood to be tested includes transmitting the light energy or propagating the light energy to the surface of the skin via the air. The output element includes a detector which receives the transmitted or reflected light energy. The output interface with the sample also includes propagating the transmitted or reflected light through the air from the skin.
Barnes et al. (U.S. Pat. No. 5,379,764) disclose a spectrographic method for analyzing glucose concentration, wherein near infrared radiation is projected on a portion of the body, the radiation including a plurality of wavelengths, followed by sensing the resulting radiation emitted from the portion of the body as affected by the absorption of the body. The method disclosed includes pretreating the resulting data to minimize influences of offset and drift to obtain an expression of the magnitude of the sensed radiation as modified.
The sensor element disclosed by Barnes et al. includes a dual conductor fiber optic probe which is placed in contact or near contact with the skin of the body. The first conductor of the dual conductor fiber optic probe acts as an input element which transmits the near infrared radiation to the skin surface while in contact therewith. The second conductor fiber of the dual conductor probe acts as an output element which transmits the reflected energy or non-absorbed energy back to a spectrum analyzer. The optical interface between the sensor element and the skin is achieved by simply contacting the skin surface with the probe, and can include transmitting the light energy through air to the skin and through air back to the probe depending upon the degree of contact between the probe and skin. Irregularities in the skin surface and at the point of measurement will affect the degree of contact.
Dahne et al. (U.S. Pat. No. 4,655,225) disclose the employment of near infrared spectroscopy for non invasively transmitting optical energy in the near infrared spectrum through a finger or earlobe of a subject. Also discussed is the use of near infrared energy diffusely reflected from deep within the tissues. Responses are derived at two different wavelengths to quantify glucose in the subject. One of the wavelengths is used to determine background absorption, while the other wavelength is used to determine glucose absorption.
The optical interface disclosed by Dahne et al. includes a sensor element having an input element which incorporates a directive light means which is transmitted through the air to the skin surface. The light energy which is transmitted or reflected from the body tissue as a measure of absorption is received by an output element. The interface for the output element includes transmitting the reflected or transmitted light energy through air to the detector elements.
Caro (U.S. Pat. No. 5,348,003) discloses the use of temporally-modulated electromagnetic energy at multiple wavelengths as the irradiating light energy. The derived wavelength dependence of the optical absorption per unit path length is compared with a calibration model to derive concentrations of an analyte in the medium.
The optical interface disclosed by Caro includes a sensor element having an input element, wherein the light energy is transmitted through a focusing means onto the skin surface. The focusing means may be near or in contact with the skin surface. The sensor element also includes an output element which includes optical collection means which may be in contact with the skin surface or near the skin surface to receive light energy which is transmitted through the tissue. Again, a portion of the light energy is propagated through air to the skin surface and back to the output element due to non-contact with the sensor and irregularities in the skin surface.
Problems with the optical interface between the tissue and the instrument have been recognized. In particular, optical interface problems associated with coupling light into and back out of the tissue were recognized by Ralf Marbach as published in a thesis entitled "MeBverfahren zur IR-spektroskopishen Blutglucose Bestimmung" (English translation "Measurement Techniques for IR Spectroscopic Blood Glucose Determination"), published in 1993.
Marbach states that the requirements of the optical accessory for measurement of the diffuse reflection of the lip are:
1) High optical "throughput" for the purpose of optimizing the S/N ratio of the spectra, PA1 2) Suppression of the insensitivity to Fresnel or specular reflection on the skin surface area.
The measurement accessory proposed by Marbach attempts to meet both requirements through the use of a hemispherical immersion lens. The lens is made out of a material which closely matches the refractive index of tissue, calcium fluoride. As stated by Marbach, the important advantages of the immersion lens for transcutaneous diffuse reflection measurements are the nearly complete matching of the refraction indices of CaF.sub.2 and skin and the successful suppression of the Fresnel reflection.
Calcium fluoride, however is not an ideal index match to tissue, having an index of 1.42, relative to that of tissue, at approximately 1.38. Thus, an index mismatch occurs at the lens to tissue interface assuming complete contact between the lens and tissue. The optical efficiency of the sampling accessory is further compromised by the fact that the lens and the tissue will not make perfect optical contact due to roughness of the tissue. The result is a significant refractive index mismatch where the light is forced to travel from the lens (N=1.42) to air (N=1.0) to tissue (N=1.38). Thus, the inherent roughness of tissue results in small air gaps between the lens and the tissue, which decrease the optical throughput of the system, and subsequently compromise the performance of the measurement accessory.
The magnitude of the problem associated with refractive index mismatch is a complicated question. First, a fraction of light, which would otherwise be available for spectroscopic analysis of blood analytes, gets reflected at the mismatch boundary and returns to the input or collection optical system without interrogating the sample. The effect is governed by the Fresnel Equation: ##EQU1## For normally incident, randomly polarized light, where N and N' are the refractive indices of the two media. Solving for the air/CaF.sub.2 interface gives an R=0.03, or a 3% reflection. This interface must be traversed twice, leading to a 6% reflected component which does not interrogate the sample. These interface mismatches are multiplicative. The fraction of light successfully entering the tissue then must be considered. In some regions of the spectrum, for instance, under a strong water band, almost all of the transmitted light gets absorbed by the tissue. The result is that this seemingly small reflected light component from the refractive index mismatch can virtually swamp out and obscure the desired signal from the sample.
Finally, it is useful to consider the critical angle effect as light attempts to exit the tissue. Tissue is highly scattering and so a light ray which launches into tissue at normal incidence may exit the tissue at a high angle of incidence. If the coupling lens is not in intimate contact with the tissue, these high angle rays will be lost to total internal reflection. The equation which defines the critical angle, or the point of total internal reflection, is as follows: ##EQU2## When light is propagating through a higher index material like tissue (N'=1.38) and approaching an interface with lower refractive index like air (N=1.0), a critical angle of total internal reflection occurs. Light approaching such an interface at greater than the critical angle will not propagate into the rarer medium (air), but will totally internally reflect back into the tissue. For the aforementioned tissue/air interface, the critical angle is 46.4.degree.. No light steeper than this angle would escape. Intimate, optical contact is therefore essential to efficient light capture from tissue.
As detailed above, each of the prior art apparatus for non-invasively measuring glucose concentration utilize a sensor element. Each sensor element includes an input element and an output element. The optical interface between the input element, output element and the skin surface of the tissue to be analyzed in each apparatus is similar. In each instance, the input light energy is transmitted through air to the surface or potentially through air due to a gap in the contact surface between the input sensor and the skin surface. Likewise, the output sensor receives transmitted or reflected light energy via transmission through air to the output sensor, or potentially through a gap between the sensor element and the skin surface even though attempts are made to place the output sensor in contact with the skin. It is believed that the optical interfaces disclosed in the prior art affect the accuracy and consistency of the data acquired utilizing the prior art methods and apparatus. Thus, the accuracy of these methods for non-invasively measuring glucose are compromised.
Accordingly, the need exists for a method and apparatus for non-invasively measuring glucose concentrations in blood which incorporates an improved optical interface. The optical interface should produce consistent repeatable results so that the analyte concentration can be accurately calculated from a model such as that disclosed by Robinson et al. The optical interface should minimize the effects on the input and output light energy due to transmission through air both into and out of the tissue being analyzed. Further, the detrimental effects of gaps due to irregularities in the surface of the skin or the presence of other contaminants should be reduced or eliminated.
The present invention addresses these needs as well as other problems associated with existing methods for non-invasively measuring glucose concentration in blood utilizing infrared spectroscopy and the optical interface associated therewith. The present invention also offers further advantages over the prior art and solves problems associated therewith.
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