Method and apparatus for the non-invasive measurement of blood glucose levels in humans

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BACKGROUND OF THE INVENTION

The present invention relates to the medical arts. It particularly relates to the measurement of a diagnostic glucose level in a human subject, especially for the monitoring of diabetic patients, and will be described with particular reference thereto. However, the invention will also find application in conjunction with the non-invasive measurement of concentrations of other proteins and other optically active substances in the human body for medical diagnosis and monitoring. For example, the invention is contemplated to be applied for measuring .beta.-amyloid protein concentrations in the body, which are indicative of Alzheimer's disease.

Diabetes is presently the fourth leading cause of mortality in the United States. Diabetes can lead to severe complications over time, including blindness, renal and cardiovascular diseases, and peripheral neuropathy associated with limbs. Diabetics typically exhibit poor blood circulation in lower extremities of the body which can lead to gangrene and subsequent amputation.

These and other diabetic complications can typically be minimized or avoided by suitable medical intervention. In the case of diabetes mellitus which relates to inadequate insulin production by the body, a regular administration of insulin injections helps convert glucose to glycogen to control diabetic symptoms and complications. The insulin-injection therapy is preferably closely monitored by frequently measuring diagnostic glucose levels. In a usual approach, blood is drawn and the serum glucose level is measured. Since this monitoring should be done regularly, e.g. on a daily basis, it is preferably self-administered, typically using a finger-prick blood extraction.

A problem arises because diabetic patients are reluctant to perform regular glucose monitoring by painful blood extraction. Blood extraction can also produce infections or introduce harmful contaminants into the body. For these and other reasons, patients sometimes neglect the invasive glucose self-monitoring and fail to adjust their insulin intake to accommodate changes and variations in glucose level. Hence, there is a continuing need for an improved and preferably non-invasive glucose monitoring method and apparatus which conveniently measures a diagnostic glucose level in the human body.

A number of approaches have been developed for determining the glucose level in ocular tissue. In particular, the glucose concentration in the aqueous humor of the eye closely mimics glucose levels in the blood. Furthermore, glucose is an optically active material whose concentration in an aqueous solution can be measured by optical polarimetric methods.

U.S. Pat. No. 5,209,231 issued to Cote et al., U.S. Pat. No. 5,560,356 issued to Peyman, and U.S. Pat. No. 6,370,407 issued to Kroeger et al. are exemplary of recent efforts to exploit the optical activity of glucose in the aqueous humor to monitor a diagnostic glucose level. However, there remains a need in the art for an improved non-invasive diagnostic glucose monitoring which is convenient for diabetic patients to use, does not require ocular implants or refractive index-matching material, provides automatic corrections for individual variations in ocular geometry and optical properties, and is optically robust and substantially insensitive to minor deviations from the designed optical alignment or configuration.

The present invention contemplates an improved apparatus and method which overcomes the aforementioned limitations and others.

SUMMARY OF THE INVENTION

According to one aspect of the invention, a method is provided for determining a diagnostic glucose level for a person. Light is reflected off an ocular lens at a Brewster's angle. A polarization rotation of the reflected light is measured after exiting the eye. A glucose concentration is determined based on the measured polarization rotation.

According to another aspect of the invention, a method is provided for determining a diagnostic glucose level. Light is reflected from an internal ocular interface at an incident angle that has a selected reflection polarization characteristic. A polarimetric parameter of the reflected light is measured. A glucose concentration is computed based on the polarimetric parameter.

According to yet another aspect of the invention, an apparatus is disclosed for determining a diagnostic glucose level in a human subject. A light source produces collimated light at a selected wavelength. The collimated light is arranged such that the collimated light passes through a portion of an eye of the subject and reflects off an eye lens at a selected angle as reflected light. A polarization analyzer measures a polarization of the reflected light that exits the eye. A path length processor determines an optical path length of the reflected light within an aqueous humor of the eye. A glucose level processor computes a glucose concentration based on the measured polarization and the determined optical path length.

One advantage of the present invention resides in providing convenient and robust non-invasive monitoring of blood glucose levels for calibrating insulin injections or other medical treatment of the diabetic condition.

Another advantage of the present invention resides in improved accuracy and precision in non-invasive measurement of glucose concentration in the human body.

Yet another advantage of the present invention resides in providing a painless method and apparatus for monitoring glucose levels in diabetic patients.

Numerous additional advantages and benefits of the present invention will become apparent to those of ordinary skill in the art upon reading the following detailed description of the preferred embodiment.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention may take form in various components and arrangements of components, and in various steps and arrangements of steps. The drawings are only for the purpose of illustrating preferred embodiments and are not to be construed as limiting the invention.

FIG. 1 schematically shows an apparatus for monitoring a diagnostic glucose level in accordance with one embodiment of the invention.

FIG. 2 shows the polarization of the exiting light for incident angles that produce light incident on the ocular lens at the Brewster's angle, at the Brewster's angle plus 0.5.degree., and at the Brewster's angle plus 1.degree..

FIG. 3 compares simulated calculations of the percentage error due to deviations from the Brewster's angle of reflection at the lens for an incident circularly polarized beam and an incident linearly polarized beam.

FIG. 4 illustrates an exemplary polarization detector which is suitable for use in the apparatus of FIG. 1.

FIG. 5 illustrates an exemplary low-coherence interferometric detector which is suitable for use in the apparatus of FIG. 1.

FIG. 6 schematically shows a fast-Fourier transform (FFT) analysis with respect to optical frequency of the output of the charge-coupled device (CCD) component of the low-coherence interferometric detector of FIG. 5.

FIG. 7 diagrammatically shows an exemplary headband head mounting of the apparatus of FIG. 1.

FIG. 8 shows an enlarged view of the substrate of the headband head mount of FIG. 7, which enlarged view diagrammatically illustrates a suitable arrangement of optical components thereon.

FIG. 9 diagrammatically illustrates an apparatus for compensating for corneal birefringence.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

With reference to FIG. 1, a glucose monitoring apparatus 10 monitors an eye 12 which includes, among other ocular tissues, a cornea 14, a lens 16, and an iris 18, which cooperate to define an anterior chamber 20 that is filled with a fluid called the aqueous humor 22. The aqueous humor includes a concentration of glucose which is to be measured.

The light-transmissive ocular tissues, in particular the cornea 14, the lens 16, and the aqueous humor 22, are each optically characterized by a refractive index denoted by "n". Suitable refractive indices include n.sub.c =1.336, n.sub.l =1.4208, and n.sub.h =1.336, for the cornea 14, the lens 16, and the aqueous humor 22, respectively. Although in this simplified model n.sub.c =n.sub.h, it is also contemplated that more precise and perhaps different refractive index values can be employed. An ambient air 24 is suitably characterized by n.sub.a =1.00.

With continuing reference to FIG. 1, the apparatus 10 includes a light source 30 that produces substantially collimated light 32. The light source 30 is preferably a multiple-wavelength light source. In FIG. 1, a multi-wavelength laser is employed. However, it is also contemplated to use other light sources that produce light at a plurality of wavelengths, such as: a white light source coupled with an optical collimator and one or more wavelength-selective filters; a mercury, sodium, or other type of arc discharge lamp; one or more light emitting diodes (LEDs); and the like. Because the light impinges upon the human eye, it should have an intensity comporting with eye safety guidelines, such as those promulgated by the American National Standards Institute (ANSI Report No. Z136.1).

The light 32 is optionally selectively polarized, e.g. using a linear polarizer 34 in combination with a quarter-wave (.lambda./4) retarder 36 to produce a circularly polarized light 38 which impinges upon the cornea 14 at an incident angle .theta..sub.I referenced to a normal 40 to the local cornea surface. (Unless otherwise noted, angles cited herein are referenced to an interface or surface normal, i.e. the normal is designated as 0.degree.).

The light 38 passes through the cornea 14 and into the aqueous humor 22 to form a transmitted light 50. The light is refracted at the tissue interfaces so that the transmitted light 50 is traveling at a transmitted angle .theta..sub.T relative to the surface of the cornea 14 through which the light passes. The angular change from .theta..sub.I to .theta..sub.T is calculable using Snell's Law according to:

n.sub.1 sin .theta..sub.1 =n.sub.2 sin .theta..sub.2 (1)

where the subscripts "1" and "2" refer to the incident and transmitted sides of the refracting interface. The transmitted light 50 impinges on the lens 16, where a portion of the light is reflected to form a reflected light 52.

The incident angle .theta..sub.I is selected such that after refraction at the ambient/cornea and cornea/aqueous humor interfaces the transmitted light 50 impinges upon the lens 16 at a Brewster's angle .theta..sub.B. Those skilled in the art will recognize that the Brewster's angle, also known as the polarizing angle, is a special angle at which the light polarization component in the plane of incidence (i.e., the p-polarization) is extinguished upon reflection, such that the reflected light is substantially polarized out of the plane of incidence, i.e., s-polarized. The s- and p-polarizations are defined with respect to the plane of incidence which contains the incident beam 50 and the reflected beam 52. The Brewster's angle for the aqueous humor/lens interface is given by: ##EQU1##

Because of the special polarizing properties of the Brewster's angle .theta..sub.B reflection, the reflected light 52 is linearly polarized, and more particularly s-polarized.

Although in the embodiment described herein a reflection at the Brewster's angle of the aqueous humor/lens interface is employed, it is also contemplated to employ reflections at other angles and/or ocular interfaces which have known polarizing properties. For example, a reflection at the lens/vitreous humor interface at the critical angle of that interface is also contemplated. Those skilled in the art recognize the critical angle as the smallest angle at which light traveling from an optically denser medium (e.g., the lens) toward a less dense medium (e.g., the vitreous humor lying behind the lens in the eye) experiences total reflection.

The reflected beam 52 traverses the aqueous humor 22. Because the aqueous humor contains substantial concentrations of glucose and other optically active substances, the polarization rotates away from the s-polarization. The rotation is given by: ##EQU2##

where .alpha..sub..lambda. is the polarization rotation at the wavelength .lambda. of the reflected light, the index i goes over all significant optically active substances including at least glucose, [.alpha.].sup.T.sub..lambda., pH,i is a specific rotation of the ith substance at the wavelength .lambda., c.sub.i is the concentration of the ith substance, and L.sub..lambda. is an optical path length. For a symmetric reflection geometry relative to a center of the lens 16, the physical path d.sub.1 of the transmitted light 50 is equal to the physical path of the reflected light 52, and the optical path length of the reflected light 52 is given by (n.sub.h.times.d.sub.1) where n.sub.h is the refractive index of the aqueous humor.

The reflected light 52 impinges upon the aqueous humor/cornea and the cornea/ambient interfaces, where the light is refracted to form exiting light 54 directed outwardly from the cornea at an exit angle .theta..sub.E relative to a normal 56 to the local cornea surface. For the symmetric geometry shown in FIG. 1, the exiting angle .theta..sub.E equals the incident angle .theta..sub.I. Asymmetric geometries respective to the center of the lens 16 can also be employed, for which the incident angle .theta..sub.I and the exiting angle .theta..sub.E are typically somewhat different.

The exiting light 54 is characterized by analyzing optics 60. A beam-splitter 62 splits the exiting light 54 into first and second beams 64, 66. The first beam 64 is analyzed by a path length detector, such as a low-coherence interferometric detector 68 which extracts the optical path length, while the second beam 66 is analyzed by a polarization analyzer or detector 70. The polarization analyzer 70 preferably extracts amplitude and phase information for both the p-polarization component and the s-polarization component of the second beam 66, e.g. in a Jones matrix or other suitable format.

The low-coherence interferometric detector 68 employs a reference beam 80 obtained from a reflected light component 82 that reflects back into the ambient 24 when the incident beam 38 impinges upon the cornea 14. The reference beam 80 is obtained using suitable optical components, such as a mirror 84 and a second beam splitter 86, to produce a combined light 88 that is analyzed by the low-coherence interferometric detector 68. For the symmetric geometry relative to the center of the lens 16 shown in FIG. 1, the low-coherence interferometric detector 68 measures an optical path length of (2.times.n.sub.h.times.d.sub.1) corresponding to a sum of the paths of the transmitted light 50 and the reflected light 52.

A variation in optical path length of about 5 percent due to eye movement and about 10 percent between individuals is expected. Measurement of the optical path length, e.g. using the low-coherence interferometric detector 68, is preferably performed to correct for such variations. However, it is also contemplated to omit the optical path length measurement and use an estimated optical path length based on the ocular geometry.

A glucose level processor 90 computes a polarization rotation .alpha. and an optical path length L.sub..lambda. from measurements of the analyzing optics 60. The exiting light 54 is related to the incident light 38 according to:

E.sub.exit =T.sub.2 T.sub.g R.sub.B T.sub.g T.sub.1 E.sub.inc (4)

where E.sub.inc is the incident beam 38 represented as a Jones vector, T.sub.1 is the Jones matrix for transmission from the ambient 24 into the aqueous humor 22 through the air/cornea/aqueous humor interface, the rightmost T.sub.g is the Jones matrix for transmission through the aqueous humor 22 from the cornea 14 to the lens 16, R.sub.B is the Jones matrix for reflection at the lens 16, the leftmost T.sub.g is the Jones matrix for transmission through the aqueous humor 22 from the lens 16 to the cornea 14 (the leftmost and rightmost T.sub.g matrices are equivalent for the symmetric reflection geometry of FIG. 1), T.sub.2 is the Jones matrix for transmission from the aqueous humor 22 to the ambient 24 through the aqueous humor/cornea/air interface, and E.sub.exit is the exiting beam 54.

For the exemplary refractive indices given previously in which n.sub.c =n.sub.h (aqueous humor having the same refractive index as the cornea), the Jones matrices are given by: ##EQU3##

where the angles .theta..sub.I, .theta..sub.B, and .theta..sub.E are as shown in FIG. 1 and .alpha..sub..lambda. is expressed in equation (3). The incident light 38 is described by a Jones vector of the form: ##EQU4##

where E.sub.0x and E.sub.0y are the amplitudes of the x- and y-components of the electric field E.sub.inc, and (.omega.t-kz) designates the spatial and temporal variation of the electric fields. The right-hand side of equation (9) is appropriate for a circularly polarized incident light.

The Brewster's or polarizing angle .theta..sub.B is expressed in equation (2) as a function of the refractive index of the aqueous humor (n.sub.h) and the lens (n.sub.1), and the incident angle .theta..sub.I can be computed using Snell's Law (equation (1)) based upon the Brewster's angle .theta..sub.B and knowledge of the ocular geometry. A suitable ocular geometric model is the Le Grand ocular model, for example as described in W. Lotmar, Journal of the Optical Society of America, volume 61, pages 1522-1529 (1971). The exiting angle .theta..sub.E can be computed similarly to .theta..sub.I.

Those skilled in the art can readily modify the expressions of equations (4)-(9) to incorporate a different incident light angle, different incident light polarization (e.g., s-polarized light), to account for a difference in refractive index between the cornea (n.sub.c) and the aqueous humor (n.sub.h), to account for refraction during transmission through the cornea 14, to account for ocular geometry variations in individual patients, and the like.

In one suitable embodiment, the glucose level processor 90 solves equation (4) based on the computed angles .theta..sub.I, .theta..sub.B, and .theta..sub.E, parameters of the selected polarization of the incident light 38, and the polarization of the exiting light 54 as measured by the polarization detector 70, to obtain the polarization rotation .alpha..sub..lambda.. However, those skilled in the art will recognize that the Brewster's angle reflection results in substantially s-polarized light as indicated in equation (7) by the zeroed p-polarization row of the Brewster's angle Jones matrix R.sub.B. This simplified optical geometry permits a simplified and more robust method for determining the polarization rotation .alpha..sub..lambda..

In a preferred embodiment, the glucose level processor 90 measures the polarization rotation .alpha..sub..lambda. as the angular shift of the major axis of the polarization ellipse of the exiting light 54 relative to the s-polarization angle which is perpendicular to the plane of incidence. Optionally, a correction is made for refractive changes due to transmission through the aqueous humor/cornea/ambient interface. The appropriate incident angle .theta..sub.I to obtain the Brewster's angle reflection at the lens 16 is suitably identified by varying the angle of incidence about the nominal value of .theta..sub.I (computed from .theta..sub.B and ocular geometric considerations) to maximize the signal-to-noise ratio. For the listed refractive indices and the exemplary ocular geometry of Le Grand, a suitable nominal angle of incidence relative to the normal to the eye is .theta..sub.I '=50.35.degree..

With continuing reference to FIG. 1, and with further reference to FIG. 2, the apparatus 10 is advantageously optically robust. In particular, the angle of incidence can vary significantly without producing substantial error in the determined polarization rotation .alpha..sub..lambda.. As seen in FIG. 2, alignment at the Brewster's angle (designated .theta.b in FIG. 2) results in a substantially linearly polarized exiting light 54. A deviation from the Brewster angle has the effect of broadening the polarization from the purely linear polarization into an elliptical polarization due to an increasingly large p-polarization component being retained upon reflection from the lens 16. However, deviations in excess of 1.degree. do not significantly change the rotational orientation of the major axis of the elliptical polarization, which is controlled by the s-polarization component due to the near-Brewster angle reflection.

With reference to FIG. 3, simulations indicate that a reduced percentage error is obtained using circularly polarized incident light as compared with incident light that is linearly polarized perpendicular to the incident plane, i.e. s-polarized. As seen in FIG. 3, deviations from the Brewster's angle of up to 4.degree. result in less than a 5% error in the glucose determination. As a result, there is a relatively large tolerance for the angular alignment of the apparatus 10, especially when circularly polarized incident light is employed.

With returning reference to FIG. 1, because the optical activity of glucose in the aqueous humor is substantially larger than that of other optically active substances, a single-wavelength glucose monitoring is contemplated, in which the glucose level processor 90 computes the glucose level from the polarization rotation .alpha..sub..lambda. according to equation (3), in which the index i runs over only a single substance, namely glucose.

Optionally, the effects of confounding optically active substances other than glucose can be corrected for by performin