WO2012134515A1

WO2012134515A1 - Method and apparatus for in vivo optical measurement of blood glucose concentration

Info

Publication number: WO2012134515A1
Application number: PCT/US2011/048968
Authority: WO
Inventors: Boris M. CHERNOBROD; Vladimir Schwartz
Original assignee: Verifica-Bg Inc.
Priority date: 2011-04-01
Filing date: 2011-08-24
Publication date: 2012-10-04

Abstract

A method of non-invasive measurement of glucose concentration directly in the blood flow by utilizing a combination of the differential scattering spectroscopy and confocal scanning laser Doppler microscopy.

Description

METHOD AND APPARATUS FOR IN VIVO OPTICAL MEASUREMENT

OF BLOOD GLUCOSE CONCENTRATION

CROSS-REFERENCES TO A RELATED APPLICATION

This application claims priority under 35 USC 119(e) to an earlier-filed U.S. Provisional patent application serial number 61/470,697, filed on April 1, 2011, which is incorporated herein in its entirety.

FIELD OF THE INVENTION

The present invention is related to the field of non-invasive optical detection and measuring of glucose concentration in blood vessels.

BACKGROUND OF THE INVENTION

It is well established that a measurement of glucose concentration directly in blood vessels is the most reliable method for monitoring of diabetic metabolism. Many methods and devices have been developed for determination of glucose in vitro or in vivo by optical means. Progress towards the development of blood glucose monitoring methods is disclosed in

"Handbook of Optical Sensing of Glucose in Biological Fluids and Tissues," edited by V.

Tuchin, CRC Press, 2009. Light scattering from the red blood cells (RBCs) is one of the noninvasive in-vivo blood glucose monitoring methods. This method exploits the fact that a change in glucose concentration leads to change in the scattering coefficient of red blood cells. A major obstacle for high accuracy measurements is parasitic scattering of light radiation in the skin tissues. Additionally, light scattering in tissues is influenced by the glucose concentration as well. Yet, the glucose concentration in tissues is not a direct manifestation of glucose

metabolism. Rather, it is a function of the local blood flow velocity, tissue temperature, oxygenation rate, etc. Additionally, other tissue constituents could influence light scattering as well. For example the glucose signal can be masked by the water absorption. Significant changes in skin tissue water concentration (+ 20%) can lead to unacceptably low accuracy and reproducibility of the blood glucose concentration measurements. Note, that the variation of water concentration in blood is relatively small (+ 1.8%). The U.S. Patent No. 5,137,023 and Published U.S. Patent Applications 2006/0063983 and 2008/0027297 disclose a technique that eliminates various influences of the skin tissues and water absorption. This invention exploits the method of differential spectroscopy by using the laser light radiation with two spectrally close wavelengths. The difference or ratio of the signals at these two wavelengths does not depend on water absorption or presence of other biological and chemical substances. This elimination of influence of the other substances is possible because the selected wavelengths correspond to sharp features in the glucose spectra. The separation of the signal from blood and the signal from tissues is based on the electronic filtering of the heart bit modulation. However, the experimental realization of this method has demonstrated unacceptably low reproducibility of the

measurements due to the influence of glucose present in skin tissues.

SUMMARY OF THE INVENTION

The present invention is directed to the method for non-invasive measuring of a blood glucose concentration, comprising directing a portion of a first laser beam and a portion of a second laser beam into a water reference cell to set a wavelength of the first laser beam to a first wavelength and to set a wavelength of the second laser beam to a second wavelength, the second wavelength corresponding to the same water absorbance as the first wavelength. The signal difference between the first wavelength and the second wavelength is sensitive to blood glucose spectral features in a refractive index or an absorption coefficient of blood; aligning the first and the second laser beams coaxially; directing the first laser beam into a measurement volume along a confocal optical path and scanning the first laser beam within the measurement volume to obtain first Doppler- shifted backscattered radiation and to obtain first-wavelength

backscattered radiation; directing the laser second beam into the measurement volume along the confocal optical path and scanning the second laser beam within the measurement volume to obtain second Doppler-shifted backscattered radiation and to obtain second wavelength backscattered radiation; angularly separating the first backscattered beam and the second backscattered beam to obtain a first separated beam and a second separated beam; directing the first separated beam to a first detector and the second separated beam to a second detector; and processing a signal from the first detector and a signal from the second detector to obtain a differential signal indicative of the blood glucose concentration.

The present invention is further directed to the method for non-invasive measuring of blood glucose concentration, comprising aligning coaxially a first laser beam and a second laser beam; directing the first laser beam into a measurement volume along a confocal optical path and scanning the first laser beam within the measurement volume to obtain first Doppler- shifted backscattered radiation and to obtain first-wavelength backscattered radiation; directing the laser second beam into the measurement volume along the confocal optical path and scanning the second laser beam within the measurement volume to obtain second Doppler-shifted

backscattered radiation and to obtain second wavelength backscattered radiation; angularly separating the first backscattered beam and the second backscattered beam to obtain a first separated beam and a second separated beam; directing the first separated beam to and the second separated beam to at least one detector; and processing a signal corresponding to the first separated beam and the signal corresponding to the second separated beam generated by at least one detector to obtain a differential signal indicative of the blood glucose concentration.

The present invention is further directed to the method for non-invasive measuring of blood glucose concentration, comprising aligning coaxially a first laser beam having a first wavelength and a second laser beam having a second wavelength, wherein the first wavelength corresponds to a higher glucose absorption and/or refractive index coefficients than that corresponding to the second wavelength; directing the first laser beam into a measurement volume along a confocal optical path and scanning the first beam within the measurement volume to obtain first Doppler-shifted backscattered radiation and to obtain first- wavelength backscattered radiation; directing the laser second beam into the measurement volume along the confocal optical path and scanning the second beam within the measurement volume to obtain second Doppler-shifted backscattered radiation and to obtain second wavelength backscattered radiation; mixing the first Doppler-shifted backscattered radiation and the first-wavelength backscattered radiation with a first reference beam having a first reference wavelength in a first interferometer and directing a first mixed signal to a first photo-detector to generate a first photo- detector signal comprised of a first alternating current and a first direct current; mixing the second Doppler-shifted backscattered radiation and the second-wavelength backscattered radiation with a second reference beam having a second reference wavelength in a second interferometer and directing a second mixed signal to a second photo-detector to generate a second photo-detector signal comprised of a second alternating current and a second direct current; and processing the first photo-detector signal and the second photo-detector signal to obtain a differential signal indicative of the blood glucose concentration.

The present invention is also directed to a blood glucose measuring device for performing the method according to any of the claims 1 through 30, comprising a source for generating a first and a second laser beams characterized by a first wavelength and a second wavelength emitting sequentially optical radiation pulses; an interferometer for receiving a first and a second reference beams split from the first and the second laser beams; an auto-focusing device for maintaining a focusing volume at a depth under a skin surface and a heating element for maintaining a predetermined skin temperature; a photodetector for detecting backscattered radiation and reference laser radiation pulses, and signals of their constructive interference; a scanning confocal optical system for the separating the backscattering radiation originating in the blood capillary vessels; and signal processing electronics for calculating a blood glucose concentration from the backscattering radiation.

The invention describes a method of non-invasive measurement of glucose concentration directly in the blood flow by utilizing a combination of the differential scattering spectroscopy and confocal scanning laser Doppler microscopy. The main sources of irreproducibility of glucose optical measurements are the influences of glucose in skin tissues and of water absorption. The differential spectroscopy method exploits the measurement of backscattering from the red blood cells (RBCs or erythrocytes) in micro-vessels by using two coaxial laser beams at two wavelengths inside a water absorption window (for example, the 2100-2400 nm window or the 1000-1560 nm window). For wavelengths inside the absorption windows, the RBC scattering has relatively sharp resonance features corresponding to the influence of the vibration resonances of glucose on mismatch of refractive index (for example, vibration resonance combination at 2300-2500 nm or first-overtone resonances at 1000-1560 nm). The laser wavelengths are near a local maximum and a local minimum of the scattering coefficient. To avoid influence of the water absorption on the extinction coefficient, the lasers wavelengths are symmetric relative to the local minimum of water absorption.

Thus, the difference or ratio of two backscattering signals is independent of the water absorption. In order to distinguish between the light scattering in tissues and blood, we apply a method of confocal scanning Doppler microscopy (CSDM). In this method, two coaxial laser beams are focused inside of the blood vessels. The backscattering beams are separated by a dispersion element such as beam splitter, dichroic mirror, grating, or Fabry-Perot resonator or other optical elements and each beam mixes with a reference beam in the interferometer. The backscattering signal includes a frequency shifted optical signal due to the scattering from moving RBC and a frequency non-shifted optical signal due to scattering from skin static structures. The interference signal oscillating at the Doppler frequency produces an alternating current (AC), at a photo-detector. The backscattering signals, including a signal at the Doppler shifted frequency, propagate through skin tissue layers such as dermis, epidermis and stratum corneum (see Figure 2). These tissue layers also contain glucose and other biochemical constituents, which affect the extinction coefficient in a way similar to the light scattering in blood.

Thus, the Doppler shifted signal acquires additional features leading to the measurement error and irreproducibility. The specificity of these features is that they are produced by the static or unmovable with respect to the laser source, structures. In order to identify this influence, we can measure a non- oscillating part of the signal responsible for a direct current (DC). By learning the dependence of the DC signal on the depth of laser beam penetration, we can measure an influence of skin tissues on AC signal. The dependence of the differential DC signal on Doppler frequency is negligibly small. Therefore, we can account for only the influence of skin tissues on the AC signal. The reference signal frequency may be shifted using a frequency shifter to eliminate a low-frequency noise. In this case, the scattering back signal from the static structures with non- shifted frequency also induces AC current of the photo-detector at a different frequency when compared with the Doppler shifted signal.

Yet another advantage of the differential confocal spectroscopy is the suppression of the signal fluctuations related to RBC motion, beam scanning over the inhomogeneous tissues and skin movements. These sources of fluctuations affect the signals at both wavelengths in the same manner due to the coaxial focusing of laser beams, which interact synchronously with the same RBC or tissue micro-volume. In other words, the fluctuations of the signals are strongly correlated in time. Thus, the relative difference or ratio of two signals is independent on these fluctuations.

Yet another aspect of the present invention is the way to non-invasively measure various blood substances and components. Such measurements are typically performed in the upper layers of human skin such as epidermis and dermis. Specifically, various blood capillaries and capillary loops in this region of interest are an object of intense studies performed by for example by laser Doppler flowmetry, optical coherence tomography, laser spectroscopy, etc. In order to ease the measurement procedure and improve its precision, thermal stimulation is often utilized to increase capillary blood flow speed and areal/volumetric density of active capillaries with flowing blood. Typical thermal stimulation cycle lasts about 10 minutes and requires maintaining of tissue temperature at approximately 40-42° C. Relatively long thermal cycle reduces the effectiveness of rapid blood test typically desired by patients and medical professionals. Due to the high areal and volumetric density of capillaries and its rapid susceptibility to thermal, electrical, reduced pressure or other appropriate methods of blood flow stimulation, the nailfold skin region appears to be as one of the most preferred regions for noninvasive measurements of various blood substances and components. It was experimentally discovered that the thermal stimulation cycle in the nailfold region at 40-42° C lasts only 15-20 seconds. During this time, the blood flow speed increases approximately 2 to 5 times from the average of 0.86 mm/sec and areal/volumetric density of active capillaries increases by 2 to 5 times as well.

BRIEF DESCRIPTION OF THE DRAWINGS

Figure 1 is a schematic illustration of measurement methodology;

Figure 2 is a schematic illustration of light-skin interaction'

Figure 3 is a schematic illustration of optimal wavelengths position relative to urea spectrum peak;

Figure 4 is an illustration of conformal surface immersion; Figure 5 is a schematic illustration of an embodiment using two interferometers and two photo- detectors;

Figure 6 is a schematic illustration of an embodiment where two lasers operate in sequence mode using one interferometer and one photo -detector;

Figure 7 is a schematic illustration of focusing inside single capillary

Figure 8 shows measurement of parasitic filtering in a static tissue above a tested capillary;

Figure 9 is an illustration of a reduced pressure device;

Figure 10 is an illustration of skin heating and reduced pressure devices;

Figure 11 shows skin heating and temperature measurement;

Figure 12 shows identification of the blood flow direction;

Figure 13 shows a possible measurement sequence;

Figure 14 shows auto-focusing, tracking, and signal reading in a nail fold region; Figure 15 shows an optical arrangement for visualization of capillary loops; Figure 16 shows a schematic illustration of a finger ring.

DETAILED DISCRIPTION OF THE INVENTION

A diagram of the measurement system is shown in Figure 1. A small portion of light from two laser beams with wavelengths ο and λι (from a tunable laser source or sources) is directed to a water reference cell. The water reference cell temperature is stabilized and is equal to the skin tissue temperature. The wavelengths of both laser beams are symmetric relative to the local minimum of water absorption window. Therefore, the water absorption is the same at both wavelengths, and the difference or ratio of corresponding backscattering signals does not depend on water absorption. Then each of the two laser beams is split into a probe beam and a reference beam. Both reference beams are directed to an interferometer. The two probe beams are arranged coaxially. To average out and reduce parasitic influence of speckle structures on backscattering signals, the probing beams are transformed into partially coherent beams by utilizing a rotating phase diffuser, deformable mirror and/or other anti-speckle devices. The partially coherent probe beams are focused inside (or under) the skin so that their beam waists are inside the blood vessels region. To average out the effects of skin' s non-homogeneity and avoid exceeding maximum permissible power exposure (MPE, described in the American National Standard for Safe Use of Lasers ANSI Z136.1), the probe beams' focusing is x-y-z scanned in three dimensions. Servo positioning and scanning of an objective lens relative to the skin surface allows selective detection of test volume within the dermis. By utilizing AC photo-detector with the signal-servo feedback loop, the region within the dermis layer with a maximum of AC photo signal is identified, and the objective lens is positioned accordingly. Backscattering beams produced by the probe beams illumination are separated by a beam splitter or by a wavelength dispersion element or elements (such as dichroic mirror, grating or a Fabry-Perot resonator, etc.). Each backscattering beam is directed to its respective interferometer or to one common interferometer. The interference signals are detected and processed. As a separator of Doppler- shifted signals, the interferometer has two additional advantages. The first advantage is the heterodyne enhancement of the signal. The second advantage is additional spatial selectivity of the signal. Only the signal waves propagating back along optical axis contribute efficiently to the interference signal. These signal waves come from the beam focus waist. Therefore, this interference suppresses parasitic background signals. These background signals include multiply scattered photons and photons produced outside of the focus volume. Both types of photons are very undesirable because they cause signal misinterpretation.

The blood glucose concentration is measured by detecting the differences in scattering (absorption) between the two probing laser beams. The refractive index of RBCs is larger than the refractive index of blood, rii > n₂ (rii is about 1.4 and n₂ is about 1.35). This refractive index mismatch causes photon scattering. Due to the movement of RBCs, this scattering is Doppler frequency shifted (this produces an AC signal, Af _max < 20 KHz). Low frequency signals due to heartbeat, vasomotions, muscle movements, etc. are filtered out electronically. Influence of water on the signal reproducibility is low because water volume concentration in blood varies by +1.8 percent (water concentration in skin tissues varies by + 20 percent). The nature of effects associated with the glucose presence and its concentration is the same for diabetics and non- diabetics. Another embodiment, shown in Fig. 6, comprises one interferometer and one photo- detector without dispersion elements for angular signal separation and uses consecutive pulses of two lasers. The time period of each laser's pulsation is much shorter than typical fluctuation times for the relevant signals in this embodiment.

Another embodiment comprises one interferometer and one laser without dispersion elements for angular signal separation and uses consecutive laser pulses at different wavelength. The time period of the laser' s pulsation is much shorter than typical fluctuation times for the relevant signals in this embodiment.

Another embodiment comprises one interferometer, two photo-detectors, and a beam splitter without dispersion elements for angular signal separation. Each reference beam, in this embodiment, interferes only with a part of a signal having the reference beam's wavelength.

Another embodiment comprises one interferometer, two photo-detectors, and a dispersion elements for angular signal beam separation. Each reference beam, in this embodiment, interferes only with a part of a signal having the reference beam wavelength.

Another embodiment, shown in Fig. 5, comprises two interferometers, two photo- detectors, and beam splitters without dispersion elements for angular signal separation. Each reference beam, in this embodiment, interferes only with a part of a signal having the reference beam's wavelength.

To reduce undesirable influence of upper skin layers, such as stratum corneum and epidermis, on the blood glucose signal, the capillary loops in close proximity to the top skin surface may be considered for obtaining reliable signal (see Fig. 7; see also Published U.S. Patent Application 2006/0142662, which is incorporated herein by reference). Using a confocal system, a signal source region (a certain volume of flowing blood) is confined in all three directions inside a capillary. Therefore, a Doppler signal may be received from the flowing blood while the influence of skin tissues is reduced.

Vasomotions of skin tissues and blood vessels require continuous adjustment to changing optimal location of the confocal test volume (see Fig. 7) using auto-focusing and auto-tracking systems based on a moving readout device with optical objective. The required accuracy is approximately +5 μιη. The optimal location of confocal test volume is established in such systems by continuous oscillating x-y-z motions of the objective in three directions.

The auto-focusing and auto-tracking systems use a Doppler backscattering signal from RBCs moving within blood vessels. The optimal Doppler signal corresponds to the optimal position of a confocal test volume relative to a capillary's position.

A differential Doppler backscattering signal can be used for focusing and tracking optimization. The advantage of using a differential signal from two lasers is absence of optical and electrical noises common to both wavelengths.

The direct measurement of parasitic filtering in the static tissues above tested capillary is achieved by scanning of the beam waist along "Z" coordinate towards the skin surface (see Fig. 8), detecting the backscattering signal from upper layers static structures, and extracting the amount of parasitic filtering by signal processing.

Due to vasomotions and movements caused by auto-focusing and auto-tracking systems, a confocal test volume may be outside a blood vessel for a period of time. Reducing the pressure directly above the skin surface increases the signal produced inside blood vessels by increasing the number of capillary loops filled with flowing blood (see Figs. 9 and 10). Reducing the pressure rapidly increases the average blood concentration and blood flow velocity in the capillaries underneath the skin.

To eliminate influence of water absorption, the laser wavelengths are symmetric relative to a local minimum of the water absorption spectrum. The position of a local minimum of water absorption depends on temperature. Therefore, the temperature of the skin tissue interacting with laser beams must be the same as the temperature of a reference water cell and the temperature of blood in micro-vessels (37 + 0.1°C). The heating device shown in Figs. 10 and 11 heats the skin test volume and measures its temperature. It comprises an optical rejection filter for wavelengths between approximately 2.1 μιη and approximately 2.3 μιη. A different set of rejection filters may protect glucose photo-detectors from the heating device radiation wavelengths. A skin- temperature measuring device is protected from lasers wavelengths by an optical rejection filter for wavelengths between approximately 2.1 μιη and approximately 2.3 μιη. To identify the direction of the blood flow inside a capillary and to differentiate arterial blood and venous blood glucose measurements, the two laser beams with different wavelengths are positioned at a rising slope and a falling slope of a capillary loop, respectively (see Fig. 12). When RBCs cross a laser beam, the backscattering signal exhibits a certain spike. The time delay between the spikes of the backscattering signals from the two laser beams determines the direction of the blood flow.

One of the possible procedures of blood glucose measurement is schematically shown in

Fig. 13.

One of the preferred places for blood glucose measurement is the nailfold region of toe or finger. Here the capillary vessels are located at close proximity to the skin surface, and have a high density and are almost horizontal in the direction of blood flow. One of the possible schemes of measurements in the nailfold region of a finger is shown in Fig. 14.

Using a CCD camera or another type of a visualizer, and using an LED as an additional light source, it is possible to choose a certain micro vessel for the measurement of blood glucose. One of the possible schemes is shown in Fig. 15.

An alternative embodiment (shown in Fig. 16) includes a finger (or another test spot) movement detection arrangement. The arrangement allows for the differentiation of the Doppler signal emanating from moving erythrocytes vs. other possible scattering sources, such as a moving skin surface structure, dense skin tissue fibers, scars, etc. Movement detection can be implemented by utilizing a MEMS gyroscope, optical and magnetic detectors or by other appropriate means.

Resonance Absorption and Anomalous Refractive Index

The com lex refractive index n is:

wherein j is the resonance number, ω is the frequency of light. . The absor tion coefficient is:

The addition art of the refractive index in case of weak dispersion is:

Note that spectral resonance features are simultaneously present in the absorption coefficient and the refractive index.

Major Blood Analytes and Their Influence on the Refractive Index

For relatively thin/shallow turbid media, the backscattering signal P_s is proportional to the reduced scattering coefficient μ_β = ¾(1— g), wherein § is the anisotropy factor: ¾ The quantity of interest is δμ , Ι μ₂. Herein, δ is the differential = _>½C¾)— _}i {¾).

The wavelength dispersion of the reduced scattering coefficient can be divided into two parts. The first part, OK_sQ, is related to the geometrical form and size of RBCs and is a universal property of Mie-like scattering. The second part, ¾< _{; i?}. appears due to the dispersion of the refractive index and depends on the contribution of analytes. The spectral resonances of glucose can contribute to the second part. Therefore, the second part depends only on the glucose concentration.

Ballistic Approximation

For relatively thin or shallow turbid media, the main contribution to the backscattering signal is by the photons that underwent only a single act of scattering. Under this assumption, we can consider the propagation of plane waves in the media with an extinction coefficient μ_£. The scattering intensity for Doppler (AC) and static (DC) signals is:

wherein ΔΩ is the spherical angle, I_L is the initial laser intensity, L_c is the beam waist length in the confocal geometry, μ_£ (z) = μ'κβο(ζ) + μ'βτ + μα is the total light scattering coefficient including the coefficients of scattering from RBCs and from static structures (ST), and the absorption coefficient μ_α. The quantity of interest is δμ'κΒο μ'ιΐΒθ, where ^RBC is the isotropic part of light scattering coefficient from red blood cells, and δμ'κβο = μ'κΒθ(λι) - μ'κΒθ(λο)·

According to the Equation (1),

where 5P_Ac = - PAC(AO)

Therefore, from the Equation (1) it follows that

Sin " " = - 2 f dzS μ Az) . (3 )

Finally,

S^s SP_{AC .} P_DC( )

=— i- ύίΠ— ί

The influence of the epidermis and stratum corneum is eliminated using the principles of confocal microscopy. The idea behind the utilization of confocal principles for direct

measurements of the blood glucose concentration is to eliminate the influence of glucose and static structures in the epidermal layer containing no blood vessels; see Fig. 2. The epidermal tissues (and the glucose contained therein) surround the blood micro-vessels and the glucose contained therein, which is measured using the scattering signal. The light scattered by the moving RBCs propagates through the skin tissues and acquires spectral features similar to those originating from the blood. The static epidermal tissues and the glucose therein act as a parasitic spectral filter for the light scattered by the RBCs. The significant difference, however, is that the signal from static structures (as compared with the signal from the RBCs) is non-reproducible and can't be utilized to assess the glucose metabolic state or its concentration.

The propagating light "remembers its history", as demonstrated by the Equation (1) having an exponential factor exp [-2 j_c i_td ]_t which is an integral over the entire optical path.

To eliminate this effect, two signals are used. The first signal is the Doppler shifted signal from the RBCs, which produces bits on the photo-detector (AC signal). The second signal is produced by the static scattering structures; it either has no Doppler shift (DC signal), or it has a different AC frequency (if the reference beam frequency is shifted by a frequency shifter). Both signals have originated from within the same test volume defined by the confocal beam waist.

Propagating along the same optical path, these two signals acquire the same parasitic spectral features. In other words, the exponential factor related to the optical path is the same for both signals.

By using an analytical combination of the measured intensity, the parasitic propagation effects are eliminated. The signal processing formula (see Equation 4) doesn't depend on the exponential factor and is free of the parasitic influence of the static scattering structures and water absorption. According to the calculations based on the Mie theory and measurements using optical coherence tomography (OCT), the major contributors to the reduced scattering at the infrared region (2200-2500 nm and 1000-1560 nm) are glucose, urea, NaCl, and KCl. The urea contribution is 7.3% of the glucose contribution. The concentration changes of NaCl and KCl are minor, and their spectral contribution is not specific in the spectral region of interest. The influence of urea on the refractive index mismatch can be reduced to 0.5% or less (from the original 7.3%) by selecting an appropriate spectral region. A selection of a spectral region is demonstrated in Fig. 3 showing the absorption spectra of glucose and urea from the data in K.V. Larin et al., "Specificity of noninvasive blood glucose sensing using optical coherent tomography: pilot study", Phys. Med. Biol., vol. 48 (2003), pages 1371-1390. The wavelengths of the lasers are chosen to be symmetric relative to the spectral peak of urea. Therefore, the difference or ratio of the signals is practically independent of the urea concentration.

Calibration Procedure

Glucose meters undergo electronic and optical calibration (as a part of the assembly procedure). Meters use the glucose calibration matrix. The data matrix is established by matching the glucose signal of calibration test stand with the glucose concentration. Glucose measurements in vivo at different levels of blood glucose concentration may be used for calibration procedure. Measurements in vitro where whole blood, plasma, or serum samples are within the hypo- to hyper-glycemic range may also be used for calibration. The glucose concentration is measured by a high precision analyzer, such as YSI 2300 STUT Plus, by YSI Life Sciences (+ 2% accuracy). The calibration test stand is based on the glucose meter design with an additional capability to test and/or to measure the specified parameters of critical performance components.

Speckle Reduction

Laser light propagating through tissues usually acquires a speckle structure due to the statistically independent scattering from various tissue structures. To avoid the undesirable signal fluctuations due to the time-dependent speckle structure, a partially coherent beam (PCB) with a time-dependent structure of coherent spots is used. A PCB is produced by transmission (or reflection) of a coherent laser beam through a light spatial phase modulator (SLM). An SLM forms a PCB by inserting into the laser phase front a time-dependent phase structure of statistically independent phase spots. A PCB produces time dependence in the speckle structure of a scattering signal emerging from the tissue. Time averaging of this time dependence by the photo-detector leads to suppression of the signal speckle structure. The PCB method permits measurements with deeper skin penetration.

An SLM may be an electro-optical phase modulator, a rotating or shifting phase diffuser, a deformable mirror, etc.

Another method of time averaging of the speckle structure is a spatial scanning of the laser beam over the tissue surface, and scanning of the laser beam's focus in a direction perpendicular to the tissue surface.

Another method of speckle structure suppression uses adaptive optics devices correcting the laser phase front to avoid appearance of speckle structures.

Time Averaging of Signal Fluctuations Produced by Non-Homogeneous Tissue Structures

Spatially non-homogeneous tissue structures, such as blood micro-vessels, fibers, and individual blood cells, may produce temporal signal fluctuations during laser beam tissue scanning. These fluctuations are suppressed by time averaging during signal processing. The averaging procedure uses an algorithm for spatial scanning, signal filtering, and modulation of the reference beam in the interferometer. Accuracy Improvement of Wavelength Tuning by Reference Cells

For high accuracy of laser wavelength tuning, thermally controlled (within + 0.1°C) reference micro-cells are used. These optically transparent micro-cells may include water, water solution of glucose, and water solution of urea and/or other blood components. The temperature control of these reference cells is in a closed-loop arrangement with a device measuring the temperature of skin tissues, such as epidermis and dermis.

Laser Wavelength Calibration

The system may contain a gas absorption cell (for example, filled with carbon monoxide) for laser wavelength verification, control, and calibration.

Skin Fresnel Reflections

The system may comprise optical components and/or arrangements to substantially reduce the effects of laser Fresnel reflections from the skin surface. These reflections reduce the measurement accuracy of the system. To further reduce the Fresnel reflections and other optical losses produced by the air/skin interface, optical surface immersion (Fig. 4) with an inflatable immersion balloon with refractive-index-matched materials (liquid and membranes) is used. The skin surface also requires a layer of immersion liquid to fill in the gaps, etc. A compliant membrane prevents optical changes in the skin tissues due to mechanical impacts, which cause expulsion of tissue water and blood from micro-vessels (and both may cause of glucose signal irreproducibility) .

Packaging

The system is hermetically sealed to avoid effects of air humidity and dust. Humidity causes parasitic absorption of laser beams and reduces the measurement accuracy. Dust causes parasitic reflections and deterioration of optical surfaces, which results in optical aberrations and reduction of measurement accuracy. Reduction of Tissue Background Signal

The system contains an arrangement (an additional set of photo-detectors, for example) to substantially reduce or suppress the effects of the background signal on the precision of the confocal signal probing.

Thermal Stabilization of Lasers, Photo-Detectors and the Opto-Mechanical Module

The system may contain the necessary arrangements and/or control loops for thermal stability of lasers and photo-detectors. Controlling thermal stability of the opto-mechanical module maintains precise positioning of optical components (such as interferometers).

Optical Focusing Arrangement

The system may comprise precise position control arrangement of the optical focus (of probing beam or beams) relative to the skin surface. This position control may comprise auto- focusing, triangulation, and/or other optical and opto-mechanical methods.

Claims

1. A method for non-invasive measuring of a blood glucose concentration, comprising: s

directing a portion of a first laser beam and a portion of a second laser beam into a water reference cell to set a wavelength of the first laser beam to a first wavelength and to set a wavelength of the second laser beam to a second wavelength, the second

wavelength corresponding to the same water absorbance as the first wavelength, wherein a signal difference between the first wavelength and the second wavelength is sensitive to blood glucose spectral features in a refractive index or an absorption coefficient of blood, aligning the first and the second laser beams coaxially;

directing the first laser beam into a measurement volume along a confocal optical path and scanning the first laser beam within the measurement volume to obtain first Doppler- shifted backscattered radiation and to obtain first-wavelength backscattered radiation;

directing the laser second beam into the measurement volume along the confocal optical path and scanning the second laser beam within the measurement volume to obtain second Doppler-shifted backscattered radiation and to obtain second wavelength backscattered radiation; angularly separating the first backscattered beam and the second backscattered beam to obtain a first separated beam and a second separated beam; directing the first separated beam to a first detector and the second separated beam to a second detector; and

processing a signal from the first detector and a signal from the second detector to obtain a differential signal indicative of the blood glucose concentration.

2. The method of Claim 1, wherein scanning of the first beam or scanning of the second beam comprises three-dimensional x-y-z scanning.

3. The method of Claim 1, wherein the step of angularly separating comprises angularly separating the first backscattered beam and the second backscattered beam by at least one spatial dispersion element.

4. The method of Claim 3, wherein at least one spatial dispersion element is a grating,

dichroic mirror, or a Fabry- Perot resonator.

5. The method of Claim 1, wherein the first and the second detectors are photodetectors.

6. A method for non-invasive measuring of a blood glucose concentration, comprising:

aligning coaxially a first laser beam and a second laser beam;

directing the laser second beam into the measurement volume along the confocal optical path and scanning the second laser beam within the measurement volume to obtain second Doppler-shifted backscattered radiation and to obtain second wavelength backscattered radiation; angularly separating the first backscattered beam and the second backscattered beam to obtain a first separated beam and a second separated beam;

directing the first separated beam to and the second separated beam to at least one detector; and

processing a signal corresponding to the first separated beam and the signal corresponding to the second separated beam generated by at least one detector to obtain a differential signal indicative of the blood glucose concentration.

7. The method of Claim 6, wherein scanning of the first beam or scanning of the second beam comprises three-dimensional x-y-z scanning.

8. The method of Claim 6, wherein the step of angularly separating comprises angularly separating the first backscattered beam and the second backscattered beam by at least one spatial dispersion element.

9. The method of Claim 8, wherein at least one spatial dispersion element is a grating, dichroic mirror, or a Fabry- Perot resonator.

10. The method of Claim 6, wherein at least one detector is at least one photodetector.

11. The method of Claim 6, comprising directing the first separated beam to the first detector and directing the second separated beam to the second detector, and processing a signal corresponding to the first separated beam generated by the first detector and the signal corresponding to the second separated beam generated by the second detector.

12. A method for non-invasive measuring of blood glucose concentration, comprising:

aligning coaxially a first laser beam having a first wavelength and a second laser beam having a second wavelength, wherein the first wavelength corresponds to a higher glucose absorbance than that corresponding to the second wavelength;

directing the first laser beam into a measurement volume along a confocal optical path and scanning the first beam within the measurement volume to obtain first Doppler- shifted backscattered radiation and to obtain first- wavelength backscattered radiation; directing the laser second beam into the measurement volume along the confocal optical path and scanning the second beam within the measurement volume to obtain second Doppler- shifted backscattered radiation and to obtain second wavelength backscattered radiation;

mixing the first Doppler- shifted backscattered radiation and the first-wavelength backscattered radiation with a first reference beam having a first reference wavelength in a first interferometer and directing a first mixed signal to a first photo-detector to generate a first photo-detector signal comprised of a first alternating current and a first direct current;

mixing the second Doppler- shifted backscattered radiation and the second- wavelength backscattered radiation with a second reference beam having a second reference wavelength in a second interferometer and directing a second mixed signal to a second photo-detector to generate a second photo-detector signal comprised of a second alternating current and a second direct current; and processing the first photo -detector signal and the second photo-detector signal to obtain a differential signal indicative of the blood glucose concentration.

13. The method of Claim 12, further comprising using a tunable laser source to generate the first laser beam and the second laser beam.

14. The method of Claim 1, further comprising using a tunable laser source to generate the first laser beam and the second laser beam.

15. The method of Claim 6, further comprising using a tunable laser source to generate the first laser beam and the second laser beam.

16. The method of Claim 1, further comprising positioning an inflatable immersion balloon having a preselected refractive index along the confocal path.

17. The method of Claim 6, further comprising positioning an inflatable immersion balloon having a preselected refractive index along the confocal path.

18. The method of Claim 12, further comprising positioning an inflatable immersion balloon having a preselected refractive index along the confocal path.

19. The method of Claim 12, wherein the first reference beam's wavelength is obtained by shifting the first wavelength with a first frequency-shifting device, and wherein the second reference wavelength is obtained by shifting the second wavelength with a second frequency-shifting device.

20. The method of Claim 12, wherein the first Doppler-shifted backscattered radiation and the second Doppler-shifted backscattered radiation originate in blood capillary loops in proximity to a skin surface.

21. The method of Claim 12, further comprising optimizing the first Doppler-shifted

backscattered radiation and the second Doppler-shifted backscattered radiation by auto- focusing and auto-tracking systems.

22. The method of Claim 21, wherein the optimizing step comprises enhancing a signal to noise ratio by positioning the measurement volume relative to a tested capillary, and further comprising enhancing the first Doppler-shifted backscattered radiation and the second Doppler-shifted backscattered radiation and improving the auto-focusing systems and the auto-tracking systems using lasers with wavelengths between 600 and 1000 nanometer.

23. The method of Claim 21, wherein the first Doppler-shifted backscattered radiation and the second Doppler-shifted backscattered radiation do not share any noise components.

24. The method of Claim 12, further comprising applying a pressure or partial vacuum

directly above a skin surface above the measurement volume.

25. The method of Claim 12, further comprising stabilizing a tempearture of the measurement volume at 37+ 0.1 °C using a heating device and a skin temperature detector.

26. A method for non-invasive measuring of blood glucose concentration, comprising:

generating a first laser beam of a first wavelength positioned at a rising slope of a capillary loop;

generating a second laser beam of a second wavelength positioned at a falling slope of the capillary loop, wherein the first wavelength corresponds to a higher glucose absorbance and/or higher contribution of glucose to the refractive index of blood than that corresponding to the second wavelength, or to a lower glucose absorbance than that of the second wavelength;

directing the first beam into a measurement volume along a confocal optical path and scanning the first beam within the measurement volume to produce first Doppler- shifted backscattered radiation and to produce first-wavelength backscattered radiation; directing the second beam into the measurement volume along the confocal optical path;

scanning the second beam within the measurement volume to produce second Doppler-shifted backscattered radiation and to produce second wavelength backscattered radiation;

mixing the first Doppler-shifted backscattered radiation and the first-wavelength backscattered radiation with a first reference beam of a first reference wavelength in a first interferometer and directing a first mixed signal to a first photo-detector to produce a first photo-detector signal combining first alternating current and first direct current;

mixing the second Doppler- shifted backscattered radiation and the second- wavelength backscattered radiation with a second reference beam of a second wavelength in a second interferometer and directing a first mixed signal to a second photo-detector to produce a second photo-detector signal combining second alternating current and second direct current; and

processing the first photo -detector signal and the second photo-detector signal to obtain a differential signal indicative of the blood glucose concentration.

27. The method of Claim 26, further comprising identifying a blood flow direction inside the capillary loop to differentiate between arterial blood and venous blood by measuring a time delay between sequential spikes of the first Doppler- shifted backscattered radiation and the second Doppler- shifted backscattered radiation.

28. The method of Claim 1, wherein method for non-invasive measuring of blood glucose concentration is performed in the nailfold region of a finger or toe.

29. The method of Claim 6, wherein method for non-invasive measuring of blood glucose concentration is performed in the nailfold region of a finger or toe.

30. The method of Claim 12, wherein method for non-invasive measuring of blood glucose concentration is performed in the nailfold region of a finger or a toe.

31. A blood glucose measuring device for performing the method according to any of the claims 1 through 30, comprising: a source for generating a first and a second laser beams characterized by a first wavelength and a second wavelength emitting sequentially optical radiation pulses; an interferometer for receiving a first and a second reference beams split from the first and the second laser beams; an auto-focusing device for maintaining a focusing volume at a depth under a skin surface and a heating element for maintaining a predetermined skin temperature; a photodetector for detecting backscattered radiation and reference laser radiation pulses, and signals of their constructive interference;

a scanning confocal optical system for the separating the backscattering radiation originating in the blood capillary vessels; and

signal processing electronics for calculating a blood glucose concentration from the backscattering radiation.