HK1111581A - Method and apparatus for monitoring glucose levels in a biological tissue - Google Patents

Description

Method and apparatus for monitoring glucose levels in biological tissue

Technical Field

The present invention relates to a method and apparatus for monitoring blood glucose levels in biological tissue. It is particularly suitable for non-invasive monitoring of blood glucose in a human or animal suffering from diabetes. In particular, it relates to methods and apparatus for non-invasive monitoring of blood glucose using optical coherence interferometry with continuous area scanning and simultaneous depth scanning to reduce the effect of speckle (noise).

Background

Monitoring blood glucose (blood glucose) concentration levels has long been critical in the treatment of diabetes in humans. Current blood glucose monitoring involves a chemical reaction of the serum with the test strip, requiring invasive blood collection by a lancet or puncture. Small hand-held monitors have been developed to enable a patient to perform this procedure anywhere, anytime. The inconvenience of this procedure-particularly blood collection and use and disposal of test strips-results in a low level of compliance. This lower compliance can lead to serious medical complications. Thus, there is a need for a non-invasive method for monitoring blood glucose.

Studies have shown that optical methods can detect small changes in light scattering of biological tissue associated with changes in blood glucose levels. Although very complex, a first order approximation of monochromatic light scattered by biological tissue can be described by the following simplified equation:

I_R＝I_O exp[-(μ_a+μ_s)L]

wherein, I_RIs the intensity of light reflected from the skin, I_OIs the intensity of light, μ, illuminating the skin_aIs the absorption coefficient, μ, of the skin for light of a particular wavelength_sIs the scattering coefficient of the skin for light of a particular wavelength, and L is the total path traversed by the light. From this relationship, it can be seen that the intensity of light decays exponentially as absorption or scattering by the tissue increases.

It has been confirmed that there is a difference between the refractive index of plasma/interstitial fluid (blood/IF) and the refractive index of the membrane of blood cells and cells such as skin cells. (see R.C. Weast et al, Handbook of CRC Chemistry and Physics, 70 th edition, (CRC Cleveland, Ohio 1989)). This difference will cause characteristic scattering of the transmitted light. Different forms of glucose are the main components of blood/IF. Changes in the glucose level in the blood/IF will change its refractive index and thus change the characteristic scattering of blood-profused tissue. In the near infrared wavelength range (NIR), blood glucose pairs change the scattering coefficient more than their absorption coefficient. Thus, the light scattering of the blood/IF and cell mixture varies with the blood glucose level. Accordingly, an optical method can be used for non-invasive measurement of the glucose concentration of blood.

Non-invasive optical techniques being investigated for blood glucose applications include polarimetry, raman spectroscopy, near infrared absorption, scattering spectroscopy, photoacoustics and optoacoustics. Despite the extensive efforts, these techniques still suffer from drawbacks such as low sensitivity, low accuracy (less than current invasive home monitors) and insufficient specificity (specificity) of glucose concentration measurements in the relevant physiological range (4-30Mm or 72-540 mg/dL). Accordingly, there is a need for an improved method for non-invasively monitoring glucose.

Optical coherence tomography, or OCT, is an optical imaging technique that uses light waves to generate high resolution images of biological tissue. OCT generates an image by interferometrically scanning a series of points in a straight line deep and measuring the absorption and/or scattering at different depths for each successive point. The data is then processed and an image of the linear section is presented. OCT has been proposed for measuring blood glucose.

However, there are serious drawbacks to using OCT for glucose monitoring. First, the OCT procedure requires long scans to reduce optical noise ("speckle"). Speckle arises from wavefront distortion when coherent light is scattered from tissue. OCT attempts to minimize speckle by averaging multiple measurements. However, this method of OCT requires a long time that is impractical for home monitors, and even so processing, speckle in OCT remains a problem for obtaining sufficiently accurate glucose level measurements.

A second disadvantage of OCT is that it requires complex processing to form an image and further processing to analyze the image data to determine glucose levels.

A third disadvantage of OCT is that it requires expensive, bulky, high precision equipment, neither suitable for transportation nor for use outside the laboratory. Accordingly, there is a need for improved methods and apparatus for non-invasive blood glucose monitoring.

Disclosure of Invention

According to the invention, the blood glucose concentration in a biological tissue is monitored by providing light having scattering properties sensitive to the glucose concentration in the tissue, and continuously scanning the light over a two-dimensional area of the tissue, while at the same time scanning the tissue in depth in a coherent measurement. Light reflected from the scanned tissue is collected and analyzed to determine the glucose concentration in the tissue. In a preferred embodiment, light from one or more light sources is split into a sample beam and a reference beam. The surface is continuously scanned with the sample beam and the phase of the reference beam is varied and used to interfere with the reflected light to achieve a depth scan in a coherent measurement mode. In a preferred embodiment, the light provided is composed of at least two different wavelengths having different scattering properties for glucose-containing tissue or for indicators of such tissue.

An apparatus for measuring blood glucose levels in biological tissue includes one or more light sources that provide light. The fiber or lens guided path directs light to the tissue, the area scanner continuously scans a two-dimensional area of the tissue with the light, and the interferometer effectively scans the tissue in depth. The interferometer also collects, analyzes and measures light reflected from within the tissue. The processor then calculates a glucose level of the blood-profused region of the tissue in response to the light measurements. Preferably, the apparatus uses a low coherence light source (light emitting diode (LED) or super luminescent diode (SLED)), a Low Coherence Interferometer (LCI), and beam focusing optics. Continuous scanning of the two-dimensional area and depth scanning can reduce noise or speckle and optimize the amount of blood-profused tissue scanned.

Drawings

The advantages, nature, and various other features of the invention will become more fully apparent upon consideration of the illustrative embodiments to which the present invention is directed. In the drawings:

figures 1a and 1b show the anatomy of the skin and the resulting OCT signal;

figures 2a and 2b illustrate the correlation between data collected from OCT signals and blood measurements and both;

FIG. 3 is a schematic diagram illustrating a blood glucose monitoring method according to the present invention;

FIG. 4 schematically illustrates an apparatus that may be used to implement the method of FIG. 3;

it should be understood that these drawings are for purposes of illustrating the concepts of the invention and are not to scale, except for the graph.

Detailed Description

Fig. 1a is a schematic cross-sectional view of the skin showing the skin surface 1, the epidermis layer 2 and the dermis layer 3. Various other structures are indicated. Certain structures, such as hairs and their accompanying hair follicles, scatter light in a manner that is independent of glucose concentration. Other areas, such as the dermal region near the capillaries, are filled with blood and interstitial fluid (blood-profused tissue), and the light scattered from these areas is highly correlated with glucose concentration.

Fig. 1b shows the scattered light as a function of depth within the skin. Area 1 of the curve corresponds to the light reflected from the skin surface. Region 2 shows light scattered within the epidermis and region 3 shows light scattered from the dermis. As we will know, the slope of the curve at the depth where the tissue is filled with blood (e.g. near the capillaries) is highly correlated with the glucose level of the blood.

Fig. 2a plots the slope of the light signal reflected by blood-profused tissue as a function of time (continuous line) versus the measured blood glucose concentration (rectangular points) after ingestion of a glucose drink by a human subject. It can be seen from the figure that the slope of the reflected signal closely follows the glucose concentration.

Fig. 2b shows a best-fit linear relationship between the slope of the reflected light signal in fig. 2a and the measured blood glucose concentration. As can be seen from the figure, the slope of the reflected light in the blood-profused tissue region is highly correlated with the glucose concentration. These data indicate a correlation of 0.95.

The applicant has determined that: the information contained in the scattered light signal can be used to measure the glucose concentration without generating and analyzing an image. At a wavelength of about 1.3 μm, the scattering coefficient μ_sIs absorption coefficient mu_aSeveral times higher and a simple linear fit to the logarithmic data will yield the scattering coefficients within the dermis layer. For blood-profused regions, the scattering coefficient can be correlated to changes in glucose concentration in the blood.

Ideally, a non-invasive glucose monitor overcomes three problems. First, speckle (noise) should be overcome within a reasonable test time. Second, the effect of those tissues other than blood-profused tissues should be minimized, and third, as much blood-profused tissue as possible should be analyzed. The applicant has solved these problems by continuously scanning a two-dimensional region of tissue to be monitored, preferably the skin, while interferometrically scanning different depths. Optionally combining area scanning with multi-slice depth scanning allows many different local areas of blood-profused tissue to be measured, minimizing the effect of speckle. Preferably, the light comprises at least two different wavelengths of light having measurably different absorption and scattering properties for glucose, blood or other biological indicators of blood-profused tissue. These two different wavelengths allow identification of which measurements are associated with blood-profused tissue.

In certain aspects, the invention can be viewed as a modification of OCT techniques: from an imaging technique to a non-imaging technique. Collection optics and data processing are greatly simplified because no image is formed. In addition, the surface scan is changed from a stepwise linear scan to a continuous or near continuous two-dimensional area scan to more quickly and efficiently reduce speckle.

Adapting the OCT system for blood glucose sensing rather than imaging provides non-invasive glucose concentration monitoring of human or animal subjects, as scattered light signal data can be more easily collected and processed, and the resulting data can be linearly correlated with glucose levels. In addition, the continuous scanning of the incident light beam over a two-dimensional area on the skin surface (rather than the conventional one-dimensional straight line) significantly improves signal stability while greatly reducing noise associated with tissue inhomogeneity. Advantageously, a region of generally circular shape is scanned.

More specifically, the present invention uses interferometry, preferably Low Coherence Interferometry (LCI), to measure glucose concentration in blood-profused tissues of humans or animals. LCI can be implemented with conventional low correlation light sources and advantageously provides a small amount of constructive interference ("optical interaction region") that can be used to localize readings to blood profused tissues. The optical interaction region can be further localized with beam focusing. LCI uses a standard interferometer illuminated by a low coherence light source. The interferometer may be any of a variety of standard forms of interferometers (e.g., michelson interferometers, mach-zehnder interferometers, etc.). The light source may be an LED or SLED.

Advantageously, the present invention is directed to measuring blood glucose concentration of tissue (3 in fig. 1a and 1 b) in the dermal layer of the skin. In this layer, the rate of light intensity decay is the scattering and absorption coefficient μ_sAnd mu_aAs a function of (c). The location of blood-profused tissue in the skin varies from subject to subject. Such tissue can be found generally in the deeper regions (greater than 600 microns) below the dermal/epidermal junction (100-350 microns deep) and below the skin surface.

With a low correlation source, an interferogram is generated for a small volume whose depth in the object can be determined by the phase of the reference beam. Which in turn may be controlled by a phase shifter, such as a movable reference mirror, preferably a mirror. Thus, a high degree of localization with respect to the measured scattering phenomena can be achieved. For example, for a typical light emitting diode operating at a wavelength of 1.3 μm, a depth resolution of 10 μm is easily achieved in biological tissue. Generally in biological tissue, scattering of light occurs at the interface of the cell membrane and the fluid surrounding the cell (e.g., blood or interstitial fluid). Measuring this scatter allows the glucose level to be determined as shown by the linear fit in fig. 2 b.

Fig. 3 is a schematic block diagram of a method of measuring blood glucose concentration in a human or animal subject. The first step, shown in block a, is to provide light having scattering or absorption characteristics that are sensitive to the glucose concentration within the tissue. Preferably, the provided light comprises at least two different wavelengths. Different wavelengths means that the wavelengths of these lights should be sufficiently different so that they have measurably different absorption and scattering properties for different levels of glucose and/or indicator components (e.g., blood). Typically, the light emitting diodes are provided by a plurality of single wavelength light sources, such as low coherence super luminescent light emitting diodes (SLEDs) having wavelengths in the infrared/near infrared range (RNIR). Alternatively, the light may be provided by a single broadband light source, suitably notch filtered.

Next, as shown in block B, the single beam of light is split into a reference beam and a sample beam. The reference beam travels along an adjustable phase path, called reference beam path (reference arm), and the sample beam travels along a sample beam path (sample arm), where the sample beam path is directed at the tissue to be monitored, such as the skin of a human diabetic. The light in the reference beam is directed onto an adjustable phase path and will then interfere with the sample light reflected from within the tissue.

In a third step, the sample beam is continuously or approximately continuously scanned over a two-dimensional region of tissue, while the depth scan is performed interferometrically, as shown in block C. Block D shows varying the phase (path length) of the reference beam so that light from the reference beam constructively interferes with reflected sample light from successively different depths of tissue. Block E shows the collection of reflected light and interference with it by the reference beam. As the interferometer scans in depth, the surface scan also continues. This can "smear" the scan and reduce the effect of speckle.

The following step (blocks F, G and H) is to process the resulting data to calculate the glucose concentration. In practice, this is achieved by calculating the scattering coefficient of glucose-containing tissue. Block F indicates that the scan data is input into the digital processor. Block G (optional but advantageous) is to determine those scatter measurements from blood-profused tissue (located in or near a blood vessel). For example, the identification may be achieved by providing light of two different wavelengths, at least one of which scatters in a characteristic manner from blood-profused tissue. Finally, in block H, the scattering coefficient of the glucose-containing tissue is calculated and the associated glucose level in the blood is determined.

Fig. 4 schematically shows an apparatus 400 for implementing the method in fig. 3. The apparatus 400 includes a fiber-based Low Coherence Interferometer (LCI). The basic interferometer is formed with a 2 x 2 fiber optic splitter 401. The optical input from the optical source 406 is split into a sample beam 402 and a reference beam 404. The sample light in beam 402 is continuously scanned across the surface of the sample by scanner 408. Preferably, the end of the sample beam 402 may include an optical imaging device 403 to adjust the size of the spot according to the area of tissue being measured. The phase of the reference beam 404 is changed or adjusted by a movable mirror 405, which can be vibrated or oscillated to perform depth scanning. The reflected signals from beams 402 and 404 interfere and are provided to photodetector 407 for measurement. Preferably, the optical imaging device 403 may provide high coupling efficiency between the optical system and the tissue.

The volume of tissue that interacts with the light, referred to as the interaction volume, is determined by the spot size (surface area) of the optical imaging device and the coherence length (depth) of the light. The reference beam 404 has a scanning mirror 405 (e.g., a mirror). The mirror 405 of the interferometer determines the phase shift applied to the reference beam 404, and therefore the reflected light from the reference beam 404 will constructively interfere with the reflected sample beam 403. The difference in phase of these beams determines the depth at which the scattered light is measured. This allows scanning at a fixed depth, an adjustable depth or multiple depths within the tissue. Thus, LCI is sensitive to the intensity of reflected light localized in a small volume of tissue. Determining the depth and interaction volume allows for more accurate selection of the area of blood-profused tissue beneath the skin.

A photodetector 407, such as a photodiode, may be used to measure the interference of light from the sample beam 402 and the reference beam 404. One or more photodetectors 407 may be used and filters may be designed for the light sources 406 of different wavelengths used in the measurements.

Preferably, the optical imaging device 403 is a beam focusing device for reducing the beam cross-section in order to minimize the optical interaction area of the tissue. The use of these optics will improve the selectivity of the signal while also mitigating the effects of speckle.

Light passing through turbid biological tissue is affected by wavefront distortion, which produces coherent noise or "speckle". By scanning different locations of the tissue multiple times and averaging the results of these scans, the effect of speckle can be mitigated. This approach is impractical for typical OCT imaging systems because the large number of scans required to reduce speckle can take too long and result in a severe loss in image resolution. However, the collection optics may become simpler for the present invention. The non-imaging system of the present invention presents a practical solution for reducing coherent noise. Not only is the effect of speckle significantly reduced, but the non-imaging system can continuously scan a two-dimensional area of tissue rather than being limited to a single scan line. Area scanning can reduce speckle due to the diversity of tissue areas involved in the scan. They also maximize coverage of blood-profused tissue. Thus, coherent noise is also further reduced.

Another approach is to use parallel light processing, in which multiple points on the subject tissue are measured together to generate a "boiling" speckle. Boiling speckle occurs when the sub-spot speckle changes so rapidly that the observed speckle is averaged out by the integration time of the human eye or light receiver. The system of the present invention can be modified to generate boiling speckle by replacing the scanner 408 with a lens array or Diffractive Optical Element (DOE). If the lens or DOE is rapidly translated or rotated along the optical axis at very high speeds, the observed speckle will be averaged out. Furthermore, the system reduces the number of scans required due to the greater diversity of speckle detected.

In the skin, since glucose is delivered to Interstitial Fluid (IF) through blood, determining the scattering coefficient in the dermal layer of tissue, which contains many blood vessels, provides the closest relevant measure of glucose concentration change. Again, the area scan increases the volume of blood-profused tissue measured.

Area scanning can be achieved by a pair of rotating prisms that continuously move the sample beam spot over a circular area of the tissue surface. Preferably, for each depth scan, the point will move a distance of a minimum of one point diameter. Thus, if the beam spot size is 12 microns and the frequency of the depth scan is 20Hz, the spot will move at a minimum rate of 240 microns per second, and preferably at a much faster rate than this.

The dots typically have a diameter of between about 10 microns and 100 microns, and preferably 20 microns and above.

The minimum area scanned is determined by the number of spot diameters required to move at the minimum depth scan rate. For a 12 micron spot and 20Hz depth scan, the minimum area that needs to be scanned is about 2200 square microns, which corresponds to a circular area of about 500 microns in diameter. More preferably, the system is designed to cover an area corresponding to a diameter of 500 to 10000 microns.

Reducing speckle using the boiling speckle method of reducing noise requires moving the multiple points fairly quickly. The spot should be moved by at least one spot diameter during the integration time of the receiver. For an individual spot size of about 10 microns and an integration time of about 4 microseconds, the spot needs to be at a minimum of 2.4 x 10⁵Speed of micron/sec.

The light source 406 can be a Light Emitting Diode (LED) or a super luminescent diode (SLED), both of which are semiconductor-based light emitters whose wavelengths can be selected to provide the best contrast between absorption and scattering of blood and other biological components, such as water. Typically, these wavelengths are in the red/near infrared (RNIR) region of the spectrum, however, longer or shorter wavelengths may also be used to increase sensitivity. Two or more light sources are advantageous for glucose measurements, and these wavelengths of light may share the same optical path through the interferometer.

A wavelength may be selected that has a minimum absorption coefficient compared to the scattering coefficients of water and blood components. If another wavelength is selected that has a peak absorption coefficient for some biological components, the difference in light attenuation of the two wavelengths may indicate the depth location of the associated structure (e.g., blood vessel). The light of these two wavelengths is absorbed by the different components in different ways. The different absorption means reduce the intensity of the scattered (reflected) light differently. The light reflected by the cell membrane is partially absorbed by the components corresponding to that wavelength. When the term "light is reflected by blood" is used, it is understood to mean that light is reflected by cells in and around the blood vessel and that certain light is absorbed by the above-mentioned components of the blood depending on the particular wavelength of the light and the glucose level of the blood. These differences in absorption and scattering properties provide the best correlation between the scattered signal and the blood glucose data.

One exemplary application is a first wavelength of about 1310nm and a second wavelength of about 820 and 960 nm. The first wavelength of 1310nm is chosen because at this wavelength the scattering properties of these liquids are at a maximum compared to the absorption properties of water, blood and blood components. The second wavelength of 820-960nm was chosen because the absorbance of light is high (compared to the first wavelength) when the blood component hemoglobin is present. If the second wavelength signal experiences a rapid decrease at a depth within the interaction volume, the rapid decrease will indicate the presence of hemoglobin and, therefore, the location of blood-profused tissue. Thus, an optimal slope region is specified for the scattering data of the first wavelength to be correlated to the glucose concentration.

A second example is where the first wavelength is about 1310nm and the second wavelength is about 1450 nm. At this second wavelength, the scattering coefficients of blood and water are similar to those at the first wavelength. However, the absorption coefficient of water at this second wavelength is exponentially higher than the absorption coefficient at the first wavelength. Thus, the different measurements of the two wavelengths indicate a change in the hydration level of the tissue. Such changes can then be used to indicate the optimal slope region for measuring blood glucose. However, the use of these two specific wavelengths also provides the additional benefit of calibrating the sensor. As the hydration level in the dermis layer changes, the scattering coefficient at the first wavelength will drift even though the glucose concentration remains stable. Thus, by measuring the hydration level of the skin using the second wavelength, this drift can be compensated for and the OCT sensor can remain in a calibrated state.

A typical method of determining glucose levels by analyzing data includes the steps of:

1. will reflect the intensity I_rAnd incident intensity I_oExpressed as logarithms, e.g. Ln (I)_r)，Ln(I_o)。

2. The logarithmic data are plotted according to the scattering formula. Because the data at the 1310nm wavelength is dominated by scattering (with minimal absorption), the logarithmic scattering formula can be expressed approximately as:

Ln(I_r)＝L_n(I_o)-(μ_t)(d)

wherein mu_tIs the scattering coefficient and d is the scan depth.

3. Determination of μ by regression analysis_t. This can be achieved by linear regression using the above mentioned logarithmic data to determine the best fit slope (μ)_t). However, because the glucose concentration is most accurately read in blood/IF, it is desirable to selectively apply regression analysis to those data points whose depth (or two-wavelength scattering property) is indicative of blood/IF. As described earlier in connection with FIG. 2b, the glucose concentration at this point is related to the scattering coefficient μ_tStrongly correlated regions and can be easily calibrated according to the scattering coefficient.

The algorithm or an equivalent algorithm can be implemented quite easily in a digital processor by those skilled in the art, providing a fast, non-invasive reading of glucose levels.

It is to be understood that the foregoing embodiments are illustrative of only a few of the many possible specific embodiments, which can represent applications of the invention. Numerous and varied other arrangements can be devised by those skilled in the art without departing from the spirit and scope of the invention.

Claims (35)

1. A method of monitoring blood glucose concentration in a biological tissue, comprising the steps of:

providing light having scattering properties sensitive to the glucose concentration within the tissue;

continuously scanning a two-dimensional region of the tissue with the light while interferometrically scanning the tissue in depth; and

collecting the reflected light to determine a glucose concentration in the tissue.

2. The method of claim 1, wherein light from one or more light sources is split into a sample beam and a reference beam,

the continuous scan comprises a scan of the sample beam; and

the interferometrically deep scan includes changing the phase of the reference beam and interfering it with the collected light.

3. The method of claim 1, wherein the light is low coherence light.

4. The method of claim 1, wherein the light comprises at least two different wavelengths having measurably different scattering properties for glucose-containing tissue or an indicator for such tissue.

5. The method of claim 1, wherein the light comprises infrared or near-infrared light.

6. A method of measuring blood glucose concentration in a human or animal subject comprising the steps of:

providing at least two different wavelengths of light, wherein each wavelength of light has a different absorption characteristic or a different scattering characteristic for glucose-containing tissue;

splitting the light into a sample beam and a reference beam;

illuminating and scanning tissue of the subject with the sample beam;

collecting sample light reflected internally from the illuminated subject tissue;

interfering with the reflected light with the reference beam; and

the interference light is measured and the measurement is processed to determine the glucose concentration in the blood of the subject tissue.

7. The method of claim 6, wherein the light is provided by one or more low coherence light sources.

8. The method of claim 6, wherein the reference light is adjusted to constructively interfere with light reflected from a selected depth within the subject tissue.

9. The method of claim 6, further comprising the step of continuously scanning tissue of the subject to reduce speckle.

10. The method of claim 6, further comprising the step of continuously scanning a two-dimensional region of the subject tissue with the sample beam.

11. The method of claim 10, further comprising the step of scanning a plurality of depths within the subject tissue to reduce speckle.

12. The method of claim 6, wherein the glucose concentration is calculated from light reflected from blood-profused tissue.

13. The method as set forth in claim 6 wherein the at least two wavelengths include wavelengths of about 1310nm and 820 nm and 960 nm.

14. The method of claim 6, wherein the at least two wavelengths comprise wavelengths of about 1310nm and 1450 nm.

15. The method of claim 6, further comprising the step of measuring changes in hydration level of the tissue.

16. The method of claim 15, further comprising the step of compensating for changes in the hydration level while glucose levels remain unchanged.

17. An apparatus for measuring blood glucose concentration of a human or animal subject, comprising:

one or more light sources for providing light at two or more wavelengths;

a first optical path for directing a sample beam from the one or more light sources to a subject tissue;

a second optical path for introducing the reference light beams from the one or more light sources to the phase shifter;

a scanner for scanning a region of the subject tissue with the sample light;

an interferometer for interfering with the reference light with sample light reflected from the subject tissue, the phase of the reference light path being adjusted to constructively interfere with the reflected sample light from a plurality of depths;

one or more photodetectors for measuring the interference light; and

a processor for calculating a glucose concentration from the measurement.

18. The apparatus of claim 17, wherein the light source is a low coherence light source and the interferometer comprises a low coherence interferometer.

19. The apparatus of claim 18, wherein the interferometer comprises:

a beam splitter for splitting light from the light source into a sample beam and a reference beam;

a sample light path for directing a sample light beam from the one or more light sources to a subject tissue;

a reference light path for guiding a reference light beam from the one or more light sources to the phase shifter; and

a light path for directing the reference beam to interfere with reflected light from the subject tissue.

20. The apparatus of claim 19, wherein the beam splitter is a 2 x 2 beam splitter.

21. The apparatus of claim 19, wherein the phase shifter is an adjustable mirror.

22. The apparatus of claim 21, wherein the mirror is adjustable to scan a plurality of depths within the subject tissue.

23. The apparatus of claim 17, wherein the photodetector comprises a photodiode.

24. The apparatus of claim 17, wherein at least one light source comprises a Light Emitting Diode (LED).

25. The apparatus of claim 17, wherein at least one light source comprises a Superluminescent Light Emitting Diode (SLED).

26. The apparatus of claim 17, wherein the scanner is a movable lens array that can be rapidly translated or rotated to reduce speckle.

27. The apparatus of claim 17, wherein the scanner is a movable diffractive optical element that can be rapidly translated or rotated to reduce speckle.

28. A method of measuring blood glucose concentration in a biological sample containing blood, the method comprising the steps of:

providing a low coherence interferometer comprising a sample beam, a reference optical path phase shifter, and one or more photodetectors;

providing a light source to provide a sample beam for illuminating the sample and a reference beam;

scanning a two-dimensional area of the illuminated sample with the sample beam;

interfering with sample light reflected from the sample with the reference beam to select light reflected within the sample blood; and

collecting and processing the interference light to calculate a measure of glucose concentration in the blood.

29. The method of claim 28, wherein the tissue region selected for sampling is blood-profused tissue.

30. The method of claim 28, further comprising the step of continuously adjusting the phase of the reference beam to scan multiple depths within the tissue.

31. The method of claim 28, further comprising the step of continuously scanning the illuminated tissue to improve the accuracy of the measurement.

32. The method of claim 31, wherein the sample and reference beams contain multiple wavelengths of light.

33. The method of claim 32, wherein the wavelength is about 1310nm and is in the range from about 820 to 960 nm.

34. The method of claim 32, wherein the wavelengths are about 1310nm and about 1450 nm.

35. The method of claim 6, wherein at least one of the two wavelengths of light is infrared or near-infrared light.