JP2003254901A - Reflection type blood sugar measuring instrument using low coherent light interferometer - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a reflection type blood sugar measuring device using a low coherent light interferometer, capable of measuring a sugar concentration from optical activity by aqueous humor in an anterior eye chamber while observing the structure of the anterior part of an eye on real time basis, with a patient's eye untouched at all. <P>SOLUTION: A light wave from a low coherent light source 1 is phase- modulated after passing through BS2, PM3, and is put in an condition wherein it is inclined from horizontality by θs by a wave length plate 4. After that, the light wave is irradiated by a lens 5 to the anterior part of an eyeball in a beam state, reflection is caused at the boundary of a crystal lens 10 and the aqueous humor 9 of the anterior eye chamber, a signal light obtained by condensing the reflected light again by the lens is propagated by a polarization fiber 11 plane maintaining, and is guided to an interferometer by a lens 12. <P>COPYRIGHT: (C)2003,JPO

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention is based on an interferometer using a low coherence light source, and the optical rotation ability of a light wave which does not reciprocate in the anterior chamber humor by the reflection type optical system in the same optical path and the anterior chamber humor. The present invention relates to a reflection-type blood glucose measuring device using a low coherence optical interferometer that non-invasively measures glucose concentration in a human body by measuring optical loss.

[0002]

2. Description of the Related Art Heretofore, there have been the following references as references in such a field.

(1) Pathophysiology Vol. 12, No. Four
(1993: 4), pp 293-300, "Noninvasive measurement method of blood glucose level-development of optical glucose sensor" (2) DIABETE S CARE, VOL. 5 N
O. 3, MAY-JUNE, 1982, pp. 259-
265 "Noninvasive Glucose
Monitoring of the Aqueous
Humor of the Eye: Part II.
Animal Studies and the Sc
(3) Crystal Physics Engineering Tomoya Ogawa, Chokabo, Applied Physics Selection, pp. 205-209 (4) APPLIED OPTICS, Vol. 37,
No. 16, 1998 June 1, pp. 3553-
3557 "Noninvasive Glucose
monitoring in vivo with
an optical heterodyne pol
Aligner ”(5) APPLIED OPTICS Vol. 39,
No. 34, 2000 December 1, pp. 6
318-6324 "Depth-resolved
two-dimensional Stokes ve
ctorsof backscattered lig
ht and Mueller matrices o
f biological tissue measure
red with optical coherence
e tomography "(6) OPTICS LETTERS, Vol. 26,
No. 13, 2001 July 1, pp. 992-9
94 "Noninvasive monitorin
go of glucose concentratio
n withdrawal coherence t
“Omography” (7) Optics Communications,
132 (1996), pp. 410-416 "Enh
anced optical rotation an
d diminished depolarizati
on in diffusive scatterin
g from a chiral liquid ”Heretofore, a method has been proposed in which an instrument is attached to the eye to measure the optical rotatory power of the transmitted light of the anterior aqueous humor. However, this method is a method in which the light wave penetrates the anterior chamber water of the eyeball, and considering actual clinical application, it is necessary to put a measuring jig in the eye,
The burden on the patient is heavy [Figs. 8 and 9 of reference (2),
See FIG. 1 in Reference (4)].

It is generally known that optical rotation is eliminated when a light wave travels back and forth through a substance having optical rotatory power. But,
In the case of a scattering medium, it is reported that the elimination of this optical rotatory power is reduced [see Reference (7)].

Currently, many patients suffering from blood glucose-related diseases, including diabetic patients, are required to develop a non-invasive measuring apparatus for internal blood glucose. On the other hand, the method using a light wave is promising because it is non-invasive, and a method using an absorption spectrum has already been reported [see Reference (1)].

Further, the anterior chamber humor in the anterior part of the eye contains relatively few substances that hinder the measurement of sugar concentration, and the sugar concentration in the anterior chamber humor is about 61% of the blood concentration. Due to its relatively good performance, its application to measurement has been attracting attention early on. For example, in 1979-1982, in
In a tro system experiment, glucose concentration was 20 to 1000 mg / 100 ml in anterior aqueous humor by optical rotation analysis using near infrared light.
It has been experimentally shown that the measurement can be performed [see References (1) and (2)].

Therefore, the inventors of the present invention have used the optical wave tomographic image measurement method (OCT: Opti) which has been practically used in ophthalmology in recent years.
cal coherence tomography)
A reflection type blood glucose measuring device using a low coherence optical interferometer, which is a basic technology of Fusing a low coherence interferometer and optical rotation analysis technology, has been proposed (Japanese Patent Application No. 2001-10575).
5 "Blood glucose meter using low coherence optical interferometer").

By the way, glucose is optically active,
Therefore, it has optical rotatory power. Optical activity means that when left / right circularly polarized light propagates through an optically active substance such as glucose,
This is a phenomenon in which the polarization state at the time of emission is different from the polarization state at the time of incidence because the refractive index for right circularly polarized light is different. As a result, since the linearly polarized light is the sum of the left and right circularly polarized light, the polarization plane of the outgoing light changes when the linearly polarized light is incident [6-3 of Reference (3)] depending on the optical activity and concentration of the substance. Figure, formula (6.25)].

Therefore, by measuring the optical rotation of this plane of polarization, the concentration is determined when the optical rotation of the substance is known. The relationship between glucose concentration and optical rotation has been experimentally measured and reported [see Reference (2)]. With a simple optical rotation analyzer, a sensitivity of 20 mg / 100 ml was obtained at 0.0013 degrees, which is a sensitivity that can be used for blood glucose measurement. Further, an experiment for determining the glucose concentration from the optical rotation in the anterior aqueous humor using heterodyne detection with a simple optical system has also been reported [see Reference (4)].

As described above, in the field of OCT, research focusing on the polarization change of living tissue has already been made.
OCT (OCT: Polarizat capable of polarization measurement)
Ion Sensitive Optical Coh
The basic structure of erence tomography) and measurement data have been reported [see Reference (5)].

From the above, there is an urgent need for research and development of a measurement method that can be measured without contacting the patient and can be expanded to a measurement system for many patients in hospitals, medical facilities, specific areas, etc. is there.

[0012]

SUMMARY OF THE INVENTION In view of the above situation, the present invention measures sugar concentration from optical activity by anterior aqueous humor while observing the structure of the anterior segment of the eye in real time without any contact with the eye of the patient. It is an object of the present invention to provide a reflection type blood glucose measuring device using a low coherence optical interferometer capable of performing the above.

[0013]

In order to achieve the above object, the present invention provides [1] a reflection type blood glucose measuring device using a low coherence optical interferometer, which comprises a low coherence light source and a low coherence light source. The light wave is phase-modulated through a beam splitter and a phase modulator, and means for obtaining irradiation light that is inclined by θs from the horizontal by a wave plate, and the irradiation light is irradiated in a beam shape to the anterior part of the eye by a lens and the anterior eye. It is characterized by comprising means for generating reflected light at the boundary with the aqueous humor and means for propagating the signal light obtained by condensing the reflected light again by the lens through the optical fiber and guiding it to the interferometer by the lens. .

[2] In the reflection type blood glucose measuring device using the low coherence light interferometer described in [1] above, since light reflected by the front surface of the crystalline lens is used, light waves are prevented from passing through the same optical path. The feature is that the optical rotation is not canceled.

[0015]

BEST MODE FOR CARRYING OUT THE INVENTION Embodiments of the present invention will be described in detail below.

FIG. 1 is a schematic diagram of a reflection type blood glucose measuring system using a low coherence optical interferometer showing an embodiment of the present invention,
FIG. 2 is an explanatory diagram of a signal light and a reference light in a reflection type blood glucose measurement system using a low coherence optical interferometer showing an embodiment of the present invention, and FIG. 3 shows a low coherence optical interferometer showing an embodiment of the present invention. It is explanatory drawing of the reflex type | mold blood glucose measuring method which was used.

As shown in these figures, this device comprises a low coherence light source (LCLS) 1, beam splitters (BS) 2 and 16, a phase modulator (PM) 3, a wave plate 4,
15, an optical fiber coupling lens 5 and 12, an optical system for observing the surface of the eyeball, and a camera unit 7. Also, 1
A stage 8 for fixing two optical fiber coupling lenses and scanning in the anterior-posterior direction of the eyeball, anterior chamber water 9, polarization plane preserving fibers 6 and 11, optical path adjusting prism 13, and mirror 1.
4, the polarizers 17 and 19, the photodetectors 18 and 20, and the differential amplifier 21.

Next, the basic principle will be described with reference to FIGS.

In FIG. 1, the light wave from LCLS1 is
Phase modulation is performed through BS2 and PM3, and the wave plate 4 is inclined by θs from the horizontal. Thereafter, the lens 5 irradiates a light wave in a beam shape on the front part of the eyeball. Since the crystalline lens 10 and the anterior aqueous humor 9 have different refractive indexes of tissues,
Reflection occurs at the boundary, and the reflected light is collected again by the lens of the camera unit 7, propagated by the polarization-maintaining fiber 11, and guided by the lens 12 to the interferometer.

As shown in FIG. 1, since the reflected light on the front surface of the crystalline lens 10 is used, the light wave does not pass through the same optical path. Therefore, the optical rotatory power does not disappear, and as shown in FIG.
The light wave is only inclined, that is, the signal light.

On the other hand, the reference optical path is matched with the signal optical path (optical path including the eyeball) and the accuracy within the coherence length of the light source by the optical path adjusting prism 13 for obtaining the interference signal. Further, due to the wave plate 15, as shown in FIG. 2A, the polarization direction becomes a light wave inclined by θr from the horizontal. The reference light and the signal light are incident on the BS 16, are separated and combined into two, and are incident on the polarizers 17 and 19 whose polarization planes are orthogonal to each other, and are detected by the photodetectors 18 and 20 for each P wave and S wave component. It is converted into an electric signal. Here, the P wave and S wave of the reference light are shown in FIG.
From (a), the following equations (1) and (2) are respectively shown.

X component (P wave) Er cos θr (1) Y component (S wave) Er sin θr (2) Further, from FIG. 2B, the P wave and S wave of the signal light are respectively expressed by (3) and equation (4).

X component (P wave) Es cos θs cos θm + Es sin θs cos θm = Es cos (θs−θm) (3) Y component (S wave) −Es cos θs sin θm + Es sin θs cos θm = Es sin ( θs−θm) (4) Here, Er and Es are the optical electric field of the reference light and the optical electric field of the signal light, respectively.

Further, since the output signals of the photodetectors 18 and 20 are given by the equations (5) and (6), the output of the differential amplifier 21 is given by the equation (7).

Output signal of the photodetector 18: I ₁ I ₁ (t) = | Er cos θr | ² + | Es cos (θs−θm) | ² + 2Er cos θr Es cos (θs−θm) cos φ (t ) = | Er | ² cos ² θr + | Es | ² cos ² (θs−θm) + 2ErEs cos θr cos (θs−θm) cos φ (t) (5) Output signal of photodetector 20: I ₂ I ₂ (T) = | Er sin θr | ² + | Es sin (θs−θm) | ² + 2Er sin θr Es sin (θs−θm) cos φ (t) = | Er | ² sin ² θr + | Es | ² sin ² (Θs-θm) + 2ErEs sin θr sin (θs-θm) cos φ (t) (6) Output signal of differential amplifier 21: I (t) I (t) = I ₁ (t) -I ₂ ( ^{t) = | Er | 2 (} cos 2 θr-sin 2 θr ^{+ | Es | 2 {cos 2} (θs-θm) -sin 2 (θs-θm)} + 2ErEs {co s θr cos (θ S -θm) -sinθr sin (θs-θ m)} cosφ (t) = | Er | ² cos ² θr + | Es | ² cos ² (θs−θm) +2 EsErcos (θr + θs−θm) cos φ (t) / ... (7) Here, the condition for sensitivity will be examined.

This is a condition that the value obtained by partially differentiating the signal output from the equation (7) with respect to θm is the maximum. The partially differentiated formula is formula (8).

I (t) = | Er | ² cos ² θr + | Es | ² cos ² (θs−θm) +2 ErEs cos (θr + θs−θm) · cos φ (t) [∂I (t)] / (∂θm) = 2 | Es | ² sin ² (θs−θm) + 2Er Es sin (θr + θs−θm) cosφ (t) ≈2ErEs sin (θr + θs−θm) cosφ (t) θr + θs−θm = π / 2 [∂I (t) ] / (∂θm) is max.

From θr, θs >> θm, θr + θs is the optimum condition.

Therefore, the condition of the analyzer is preferably θr = θs = π / 4. (8) From this, θr + θs−θm = π / 2 is the highest sensitivity condition, and θm is minute, so eventually θr + θs = π / 2.

Under this condition, the output signal is given by equation (9).

When θr = θs = π / 4, cos2 θr = cos2 (π / 4) = cos (π / 2) = 0 cos2 (θs−θm) = cos (2θs−2θm) = cos2θs · cos2θm + sin2θs · sin2θm = sin2θm cos (θr + θs−θm) = cos (π / 2−θm) = cosπ / 2 cosθm + sinπ / 2 · sin θm = sin θm ∴I (t) = | Es | ² · sin2θm + 2EsEr · sin θm · cos φ (θ) = where θr = = Π / 4) (9) Further, since θm is minute, approximately θm = sin θm
And the signal light is weaker than the reference light, so Er
>> Es holds. At this time, the output signal is given by the equation (10).
Becomes

In general, θm≈sin θm, Er >> Es
Therefore, ∴I (t) = 2EsErθm · cosφ (t) (10) In this equation, since the optical electric field Er of the reference light is sufficiently large, the product Er · Es can be sufficiently detected even when the optical electric field Es of the signal light is small. That is, the sensitivity is increased by heterodyne detection.

Next, the measuring method will be described with reference to FIG.

First, the coherence gate is aligned with the intersecting region of the two beams using the prism 13 (see FIG. 1) for adjusting the optical path. When the signal output is recorded while moving the lens stage 8 so that the intersecting region is aligned with the surface of the cornea and bringing the lens stage 8 close to the eyeball, the signal change shown in FIG. 3, that is, the reflected light intensity from the front surface of the cornea. S ₁ ,
The reflected light intensity S ₂ from the front surface of the lens and the reflected light intensity S ₃ from the rear surface of the lens are measured. At this time, the optical loss in the anterior aqueous humor 9 can be quantitatively evaluated from the ratio of S ₁ and S ₂ , and information completely independent of the glucose concentration can be obtained. In Reference (6), a method for obtaining the glucose concentration from the optical loss is reported, which is required completely different from the method utilizing the optical activity (the method proposed here is not only optical activity but also optical activity). It is also possible to measure glucose concentration using loss.It is important to improve the measurement accuracy by simultaneous measurement using multiple methods in order to reduce errors due to individual differences and differences in measurement conditions.)

Considering the signal change of FIG. 3, S ₁ , S ₂
Θm is obtained from the change in the signal output of the differential amplifier 21 by adjusting the stage 8 of the lens to the position where the above is obtained.
As a method of obtaining the glucose concentration C from θm, generally, the following relation C as shown in the reference (4) is used.
= Θm / (αL) holds. Here, C: glucose concentration, θm: optical rotation angle, α: glucose substance constant, L:
Propagation distance.

From the above, L is obtained from the interval between S ₁ and S ₂ , so C is obtained.

In an actual application, it is more systematic and practical if a main measuring instrument is provided in a central measuring room in a hospital or the like, and a probe can be distributed to each patient by an optical fiber for measurement.

A measurement system that meets such needs will be described below.

FIG. 4 is a schematic diagram of a reflection type blood glucose measuring system using a low coherence optical interferometer showing another embodiment of the present invention, and FIG. 5 shows a low coherence optical interferometer showing another embodiment of the present invention. It is explanatory drawing of each light wave in the reflection-type blood glucose measurement system used.

As shown in FIG. 4 (a), this apparatus has a low coherence light source (LCLS) 25, a wave plate 26, beam splitters (BS) 27, 30, an optical path channel switch 28 for selecting a probe for a patient, Polarization maintaining fiber 29, phase modulator (PM) 31, mirrors 32, 3
3, a polarization beam splitter (PBS) 34, photodetectors 35 and 36, and a differential amplifier 37.

The optical probe for each patient is shown in FIG.
As shown in (b), a polarization-maintaining fiber 29, a polarization-maintaining coupler 38, a half mirror 39 that partially reflects a light wave, an optical isolator (OI) 42 that makes the propagation direction of the light wave one direction, and a coupling lens 43. , A coupling lens stage 44.

The basic principle will be described with reference to FIGS. 4 and 5.

In FIG. 4A, the light wave from the LCLS 25 is guided to the optical probe (polarization plane preserving fiber 29) through the wave plate 26, the BS 27, and the optical path channel switch 28 for selecting the patient probe. .

In the optical probe (polarization maintaining fiber 29) shown in FIG.
Although it is branched into one fiber, since it is unidirectional by the optical isolator 42, propagation to the other fiber does not occur, but only clockwise propagation.

The half mirror 39 reflects the reflected light (reference light) 4
0 is generated, and this becomes the reference light. The transmitted light 41 is applied to the front surface of the crystalline lens by the coupling lens 43-1 and the reflected light accompanied by the optical rotation becomes signal light, which is introduced into the optical fiber by the coupling lens 43-2 and returns to the interferometer through the polarization-preserving coupler 38. As described above, in the optical probe (polarization-maintaining fiber 29), the reference light with a delay is generated due to the difference between the signal light accompanied with the optical rotation and the optical path.

The two light waves enter the Michelson interferometer via the optical path channel switch 28 and BS 27 which selects the probe for the patient. Both light waves are separated into two by the BS 30, one of which passes through the phase modulator 31 for heterodyne detection, is reflected by the mirror 32, and is reflected by the PBS 34.
Incident on.

The other side is reflected by the mirror 33, and PB
Go to S34. The mirror 32 gives an optical path difference of L in order to compensate for the delay between the signal light and the reference light. PBS
In 34, the P wave component is detected by the photodetector 35, the S wave component is detected by the photodetector 36, and the differential signal component is output by the differential amplifier 37.

The delay compensation will be described with reference to FIG.

FIG. 5A shows a light wave from the optical probe [reflected light (reflected light) 46 from the half mirror and reflected light (signal light) 47 on the front surface of the lens rotated]. The reference light 40 generated by the half mirror 39 has a shorter propagation optical path than the signal light that is reflected and rotatively reflected on the surface of the crystalline lens, and therefore returns to the interferometer earlier than the signal light 49 by t ₀ . Since both light waves are deviated in time, it is impossible for both to interfere with each other. When an optical path difference was given to the Michelson interferometer, the light wave immediately before the PBS 34 was delayed again as shown in FIG. 5B, the reference light 48 and the signal light 49, and as shown in FIG. 5C. The signal light 51 is delayed from the reference light 50. Those light waves are L
= C · t ₀ holds, the signal light 49 and the reference light 50 coincide with each other in time, as shown in FIG. 5, so that interference occurs. These are separated and detected for each P wave and S wave by the PBS 34 and become the output signal of the differential amplifier 37. Here, t ₀ represents the delay due to the optical path difference L.

The present invention is not limited to the above embodiments, and various modifications can be made within the scope of the present invention, and these modifications are not excluded from the scope of the present invention.

[0051]

As described in detail above, according to the present invention, the following effects can be achieved.

(A) Since the sugar concentration can be measured from the optical activity in the anterior aqueous humor while observing the structure of the anterior segment of the eye in real time without any contact with the patient's eye, the burden on the patient is reduced. , The measurement time is fast. Moreover, since it is possible to support systemization using an optical network, it is possible to provide services to patients in many remote areas.

(B) According to the present invention, it is possible to measure blood glucose with a small burden on many diabetic patients, and therefore, the contribution to the medical field and the ripple effect to the semiconductor and other industrial fields are great. is there.

[Brief description of drawings]

FIG. 1 is a schematic view of a reflective blood glucose measurement system using a low coherence optical interferometer showing an embodiment of the present invention.

FIG. 2 is an explanatory diagram of signal light and reference light in a reflective blood glucose measurement system using a low coherence optical interferometer showing an embodiment of the present invention.

FIG. 3 is an explanatory diagram of a reflective blood glucose measurement method using a low coherence optical interferometer showing an embodiment of the present invention.

FIG. 4 is a schematic diagram of a reflective blood glucose measurement system using a low coherence optical interferometer showing another embodiment of the present invention.

FIG. 5 is an explanatory diagram of each light wave in a reflection-type blood glucose measurement system using a low coherence optical interferometer showing another embodiment of the present invention.

[Explanation of symbols]

1,25 Low coherence light source (LCLS) 2,16,27,30 Beam splitter (BS) 3,31 Phase modulator (PM) 4,15,26 Wave plate 5,12 Optical fiber coupling lens 6,11,29 Polarization maintaining fiber 7 camera unit 8 stages 9 Anterior aqueous humor 13 Prism for optical path adjustment 14,32,33 mirror 17,19 Polarizer 18, 20, 35, 36 Photodetector 21,37 Differential amplifier 28 Optical path channel switch 34 Polarizing Beam Splitter (PBS) 38 Polarization preserving coupler 39 Half mirror 40 Reflected light (reference light) 41 transmitted light 42 Optical Isolator (OI) 43-1 and 43-2 coupling lens 44 coupled lens stage 46,48,50 reference light 47,49,51 Signal light

   ─────────────────────────────────────────────────── ─── Continued front page    F term (reference) 2G059 AA01 BB12 CC16 EE02 EE05                       EE09 GG04 GG06 GG10 JJ11                       JJ12 JJ17 JJ18 JJ20 KK01                       KK03 KK04                 4C038 KK10 KL05 KL07 KX04

Claims (2)

[Claims] 1. A low coherence light source (a); and (b) a light beam from the low coherence light source, a beam splitter,
A means for obtaining irradiation light that is inclined by θs from the horizontal by a wave plate while performing phase modulation through a phase modulator, and (c) irradiating the irradiation light in a beam form on the anterior part of the eye by a lens,
A means for generating reflected light at the boundary between the crystalline lens and the anterior aqueous humor, and (d) means for propagating the signal light obtained by condensing the reflected light again by the lens through the optical fiber and guiding it to the interferometer by the lens. A reflection-type blood glucose measuring device using a low-coherence optical interferometer, which is characterized by being provided. 2. The reflection-type blood glucose measuring apparatus using the low coherence optical interferometer according to claim 1, wherein the reflected light on the front surface of the crystalline lens is used, so that the light waves do not pass through the same optical path, and the optical activity is reduced. A reflection-type blood glucose measuring device using a low-coherence optical interferometer, which is characterized in that the above-mentioned problem is prevented.