VIVO CALIBRATION OF DOPPLER FLOWMETRY

Abstract

A method for determining a calibration factor in Doppler flowmetry velocity measurements in the living eye includes imaging the eye with Doppler flowmetry and processing data to obtain blood velocity, volume, and flow maps using Doppler flowmetry formulas that provide velocity as a mean frequency expressed in Hz, and volume and flow in arbitrary units. A selected blood vessel is probed with Doppler OCT to measure the absolute velocity of blood at that location expressed in mm/s to determine a calibration factor used to convert the velocity measured with Doppler flowmetry expressed in Hz to velocity expressed in mm/s.

Claims

1. A method for determining a calibration factor in Doppler flowmetry velocity measurements in the living eye, the method comprising: imaging the eye with Doppler flowmetry; processing the data to obtain blood velocity, volume, and flow maps using Doppler flowmetry formulas that provide velocity as a mean frequency expressed in Hz, and volume and flow in arbitrary units; probing a selected blood vessel with Doppler OCT to measure the absolute velocity of blood at that location expressed in mm/s; and determining a calibration factor to convert the velocity measured with Doppler flowmetry expressed in Hz to velocity expressed in mm/s.

2. The method in claim 1 wherein the living eye is a human eye or an animal eye.

3. The method in claim 1 wherein the Doppler flowmetry method is line-scanning Doppler flowmetry (LSDF).

4. The method in claim 1 wherein the calibration factor is calculated using LSDF and OCT measurements at the same location in the retina and at the same time.

5. The method of claim 1 in which probing a selected blood vessel with Doppler OCT includes: scanning the OCT beam in a circular pattern; intersecting the blood vessel twice, identified by two spots in the OCT image; measuring the height difference of the two spots; calculating the angle between the blood vessel and the OCT beam using the measured height difference; and calculating the absolute blood velocity using the angle between the blood vessel and the OCT beam and the axial velocity component obtained from Doppler OCT.

6. The method in claim 5 wherein the circular OCT scan is generated by scanning an OCT beam on the surface of a cone with the apex in the center of the eye pupil.

7. The method in claim 6 wherein the cone angle is approximately 2 degrees.

8. The method of claim 1 in which obtaining blood volume includes: fitting a profile of the Doppler flowmetry data in a plane perpendicular to the blood vessels with a parabolic function; and calculating the diameter of the blood vessel as the distance between the two points where the parabolic fit function is zero.

9. The method in claim 8 wherein the Doppler flowmetry data is a velocity map.

10. The method in claim 8 wherein the Doppler flowmetry data is a volume map.

11. The method in claim 8 wherein the Doppler flowmetry data is a flow map.

12. The method of claim 8 in which calculating the cross-sectional area of the blood vessels includes using a calculated blood vessel diameter.

13. The method of claim 12 in which obtaining the volumetric flow map expressed in mm.sup.3/s, comprises multiplying the calibrated velocity map by the calculated cross-sectional area map.

14. A method of determining blood velocity measurements in the living eye, the method comprising: imaging the eye with Doppler flowmetry; obtaining blood velocity, volume, and flow maps using Doppler flowmetry formulas that provide velocity as a mean frequency expressed in Hz, and volume and flow in arbitrary units; probing a selected blood vessel with Doppler OCT to measure the absolute velocity of blood at that location expressed in distance over time; determining a calibration factor to convert the blood velocity measured with Doppler flowmetry expressed in Hz to blood velocity expressed in distance over time; and calculating the blood velocity in distance over time using the calibration factor.

15. A method for determining blood velocity measurements in the living eye, the method comprising: imaging the eye with Doppler flowmetry; obtaining velocity, volume, and flow maps using Doppler flowmetry formulas that provide blood velocity as a mean frequency expressed in Hz and volume and flow in arbitrary units; probing a selected blood vessel with Doppler OCT to measure the absolute velocity of blood at that location expressed in distance over time; determining a calibration factor to convert the blood velocity measured with Doppler flowmetry expressed in Hz to velocity expressed in mm/s; and using the calibration factor to convert the blood velocity measured with Doppler flometry to blood velocity expressed in distance over time.

16. The method in claim 15 wherein the living eye is a human eye or an animal eye.

17. The method in claim 15 wherein the Doppler flowmetry method is line-scanning Doppler flowmetry (LSDF).

18. The method in claim 15 wherein the calibration factor is calculated using LSDF and OCT measurements at the same location in the retina and at the same time.

19. The method of claim 15 in which probing a selected blood vessel with Doppler OCT includes: scanning the OCT beam in a circular pattern; intersecting the blood vessel twice, identified by two spots in the OCT image; measuring the height difference of the two spots; calculating the angle between the blood vessel and the OCT beam using the measured height difference; and calculating the absolute blood velocity using the angle between the blood vessel and the OCT beam and the axial velocity component obtained from Doppler OCT.

20. The method in claim 18 wherein the circular OCT scan is generated by scanning an OCT beam on the surface of a cone with the apex in the center of the eye pupil.

21. The method in claim 19 wherein the cone angle is approximately 2 degrees.

22. The method of claim 15 in which obtaining blood volume includes: fitting a profile of the Doppler flowmetry data in a plane perpendicular to the blood vessels with a parabolic function; and calculating the diameter of the blood vessel as the distance between the two points where the parabolic fit function is zero.

23. The method in claim 2 2 wherein the Doppler flowmetry data is a velocity map.

24. The method in claim 22 wherein the Doppler flowmetry data is a volume map.

25. The method in claim 22 wherein the Doppler flowmetry data is a flow map.

26. The method of claim 22 in which calculating the cross-sectional area of the blood vessels includes using a calculated blood vessel diameter.

27. The method of claim 25 in which obtaining the volumetric flow map comprises multiplying the calibrated velocity map by the calculated cross-sectional area map.

28. A method for determining a calibration factor in Doppler flowmetry velocity measurements in the living eye, the method comprising: imaging the eye with Doppler flowmetry; processing the data to obtain blood velocity, volume, and flow maps using Doppler flowmetry formulas that provide velocity as a mean frequency expressed in Hz, and volume and flow in arbitrary units; probing a selected blood vessel with Doppler OCT to measure the absolute velocity of blood at that location expressed in distance over time by: scanning the OCT beam in a circular pattern, intersecting the blood vessel twice, identified by two spots in the OCT image, measuring the height difference of the two spots, calculating the angle between the blood vessel and the OCT beam using the measured height difference, and calculating the absolute blood velocity using the angle between the blood vessel and the OCT beam and the axial velocity component obtained from Doppler OCT; and determining a calibration factor to convert the velocity measured with Doppler flowmetry expressed in Hz to velocity expressed in distance over time.

29. The method in claim 28 wherein the living eye is a human eye or an animal eye.

30. The method in claim 28 wherein the Doppler flowmetry method is line-scanning Doppler flowmetry (LSDF).

31. The method in claim 28 wherein the calibration factor is calculated using LSDF and OCT measurements at the same location in the retina and at the same time.

32. The method in claim 28 wherein the circular OCT scan is generated by scanning an OCT beam on the surface of a cone with the apex in the center of the eye pupil.

33. The method in claim 32 wherein the cone angle is approximately 2 degrees.

34. The method of claim 28 in which obtaining blood volume includes: fitting a profile of the Doppler flowmetry data in a plane perpendicular to the blood vessels with a parabolic function; and calculating the diameter of the blood vessel as the distance between the two points where the parabolic fit function is zero.

35. The method in claim 34 wherein the Doppler flowmetry data is a velocity map.

36. The method in claim 34 wherein the Doppler flowmetry data is a volume map.

37. The method in claim 34 wherein the Doppler flowmetry data is a flow map.

Description

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWINGS

[0030] Other objects, features and advantages will occur to those skilled in the art from the following description of a preferred embodiment and the accompanying drawings, in which:

[0031] FIG. 1 shows an ultra-widefield velocity montage of the left eye of a normal volunteer illustrating visualization of vortex veins without injecting contrast agent;

[0032] FIG. 2 illustrates a LSDF velocity map with potential locations of the OCT probing beam;

[0033] FIG. 3 is a diagram of the scanning geometry;

[0034] FIG. 4 shows examples of OCT images of a glass tube at different tilt angles embedded in scattering medium;

[0035] FIGS. 5A-B show examples of LSDF images (SLO-left and Velocity-right) illustrating the flow in the glass tube and the location of the OCT circular scan;

[0036] FIGS. 6A-B show the parabolic fit of the un-wrapped phase images;

[0037] FIG. 7 shows the results of tube diameter measurement with OCT;

[0038] FIG. 8 shows the results of angle measurement with OCT;

[0039] FIG. 9 shows the results of velocity measurement with OCT;

[0040] FIG. 10 shows the results of the volumetric flow rate measurement with OCT;

[0041] FIG. 11 shows the results of Velocity measured with LSDF;

[0042] FIGS. 12A-B illustrate various LSDF power spectra across the flow tube;

[0043] FIG. 13 shows changes of the power spectrum shape with increasing flow velocity;

[0044] FIG. 14 shows a fit of the LSDF measured Velocity with an error function;

[0045] FIG. 15 shows the calibrated LSDF Velocity;

[0046] FIGS. 16A-B illustrate a fit of LSDF measured Volume (A) and the tube diameter measurements (B); FIG. 17 shows the LSDF measured flow rate as a function of the set flow rate;

[0047] FIG. 18 shows the LSDF measured flow rate as a function of the OCT measured flow rate;

[0048] FIG. 19 shows examples of OCT reflectivity and phase images and the corresponding scan diagram in the small circular scan configuration;

[0049] FIG. 20 illustrates the LSDF Velocity image with the location of the OCT circular scan and the corresponding OCT images with the two spots automatically identified by the processing software;

[0050] FIG. 21 is a block diagram depicting in one example, the primary components associated with a combined LSDF/OCT system;

[0051] FIG. 22 is a diagram of the optical components of the system of FIG. 21; and

[0052] FIGS. 23A and B are schematic views of an exemplary LSDF/OCT instrument.

DETAILED DESCRIPTION OF THE INVENTION

[0053] Aside from the preferred embodiment or embodiments disclosed below, this invention is capable of other embodiments and of being practiced or being carried out in various ways. Thus, it is to be understood that the invention is not limited in its application to the details of construction and the arrangements of components set forth in the following description or illustrated in the drawings. If only one embodiment is described herein, the claims hereof are not to be limited to that embodiment. Moreover, the claims hereof are not to be read restrictively unless there is clear and convincing evidence manifesting a certain exclusion, restriction, or disclaimer.

[0054] In retinal blood flow Doppler (LSDF) imaging, the measurements are given in arbitrary units (Volume and Flow) or Hz (Velocity) and their conversion factor to the correct parameters claimed to be measured were rather elusive so far. The measurements seem to depend on a lot of uncontrolled factors, are operator dependent among other things, and cannot reliably support longitudinal studies on the same eye or comparisons among individuals.

[0055] Disclosed here is a method to calibrate the LSDF velocity map by determining the calibration factor from Hz to mm/s based on a local OCT measurement as illustrated in FIG. 2. Potential OCT probe beam circular scan locations are shown as small white circles.

[0056] Before performing measurements on human volunteers, we first tested the ability of the OCT system to quantify the flow velocity and the geometric dimensions of the flow channel, as described below. We used a cylindrical glass tube embedded in a scattering medium for a realistic testing of the quantitative flow measurement in a configuration that resembles the flow in a human eye. The flow channel was mounted on a rotation stage that allowed for controlled orientation of the flow with respect to the laser beam and on a micrometric translation stage for placing it in the focal plane of a 25 mm focal length lens. This arrangement acts as a model eye with the lens simulating the eye lens. Measurements were also made holding the flow channel stationary and changing the offset voltage of the scan to change the orientation of the scan vertically as one would do in a real eye to reposition the scan on a blood vessel. The liquid flown through the microfluidic channel was milk diluted in water.

[0057] The inside diameter of the cylindrical glass tube was 198 μm. The tube was connected to a syringe pump using silicone tubing to which was glued with epoxy. For a very stable phantom that could be used repeatedly over longer periods of time we used 3 μm aluminum oxide powder mixed in transparent silicone sealant. The scanning geometry is shown in FIG. 3. The OCT beam was scanned over the surface of a cone with the pivot point A. The scan intersects the flow twice and the height difference of the two spots is used to calculate the angle between the flow and the laser beam.

[0058] FIG. 4 shows examples of the OCT images from the same tube with different tilt angles. LSDF measurements were also performed and an example is shown in FIG. 5 where the white circle illustrates the position of the OCT circular scan with respect to the tube.

[0059] The dark ring around the spots in FIG. 4 is the glass wall of the tube which is clear without any scattering particles. Doppler OCT calculation within the two white spots provides the Doppler phase shift AO (in radians) due to the flow of the diluted milk through the tube. This phase shift provides the axial component of the flow velocity, along the OCT beam. The vertical position of the two spots is used to calculate the tilt angle θ between the flow (tube) and the OCT beam which is then used to calculate the absolute velocity of the flow in each pixel using the following formula:

[00001]$\begin{array}{cc}V=\frac{\lambda }{4\pi n\Delta T}\frac{\mathrm{\Delta \varnothing }}{\cos \theta }& \left(1\right)\end{array}$

where: λ=1.06 μm is the OCT central wavelength, n=1.33 is the refractive index of water, and ΔT= 1/70 ms is the time between consecutive A-lines (for 70 kHz A-line rate).

[0060] 200 circular OCT scans were acquired and the phase calculations were averaged over these 200 scans assuming constant velocity flow during the data acquisition time (˜2.86 s). The total angle of the cone scanning geometry was approximately 2° (1° half angle).

[0061] One issue with Doppler OCT is that the phase calculation involves an arctangent which wraps at ±π/2. As the velocity and the angle increase, the phase approaches π/2 and then it jumps down to −π/2 resulting in alternating positive and negative rings. A useful procedure that provides rapid automatic un-wrapping of 2D phase images is based on least-squares, iteration and calibration to phase derivatives is described in Xia, H., et al., Phase calibration unwrapping algorithm for phase data corrupted by strong decorrelation speckle noise, Opt. Express 2016. 24(25): p. 28713-28730 and Xia, H., et al., Non-invasive Mechanical Measurement for Transparent Objects by Digital Holographic Interferometry Based on Iterative Least-Squares Phase Unwrapping, Experimental Mechanics, 2012. 52(4): p. 439-445 incorporated herein by this reference.

[0062] The OCT reflectivity images shown in FIG. 4 allow for segmentation of the tube position and calculation of the vertical location of each of the two spots. The parabolic fit of the Doppler phase shift shown in FIG. 6 also enables localization of the center of the flow. Either one can be used for angle calculation.

[0063] FIG. 6B shows the parabolic fit in the vertical (axial) direction which enables calculation of the tube diameter D from the two positions where the parabola gets to zero as there is no flow at the tube wall. A few pixels on both sides of the parabola were removed because they are affected by intensity and phase artifacts due to reflections at the liquid/glass interface.

[0064] FIG. 7 shows the results of tube diameter measurement with OCT and the three orientations over the entire range of set velocities up to 130 mm/s. The x symbol was used for 3° tilt, the + sign for 6° and the o symbol for 9° tilt. The line 20 is set for the actual value of the tube diameter as provided by the manufacturer with an uncertainty of ±5 μm. It should be noted that most of the measurements are within ±5% of the expected value with the exception of low velocity values below 10 mm/s where the errors are larger. The results are slightly less precise for 3° which is expected since the Doppler shift calculation is most affected by phase noise when the beam is close to normal to the flow.

[0065] The measurement of the absolute velocity is based on Eq. 1 which requires estimation of the angle θ between the flow and the laser beam. The results for the angle measurement are shown in FIG. 8. As mentioned above, the flow channel was mounted on a rotation stage that allowed for controlled orientation of the flow with respect to the laser beam. The intended angle between the axis of the scanning cone and the tube was set at 3°, 6°, and 9°, as it is called out in all the plots (3°-x, 6°-+, 9°-o). However, the set rotation was with respect to the vertical axis and the tube was not always orthogonal to this axis. There might have been slight misalignments between the tube and the slide holding the tube and between the slide and the rotation stage and/or the axis of the scanning cone might not have been always perfectly horizontal. Therefore, some additional compound angles which were not easy to measure slightly affected the actual angle between the tube and the beam.

[0066] The line for each measurement set in FIG. 8 is a fit through the data and therefore an estimate of the actual angle between the flow tube and the OCT beam. Most of the times it is close to the set value, but not exact. However, the measured value is remarkably constant across the entire set velocity range (with some exceptions below 5 mm/s). This angle value was used for velocity calculation shown in FIG. 9.

[0067] The results of the velocity measurement from OCT data are shown in FIG. 9 with the same colors and symbols as mentioned above. There is also a line 22 in that plot, generally buried under data, as a guide for one-to-one correspondence between the measured (vertical) and set (horizontal) values. The experimental values are remarkably close to this guide with small errors noted again for 3° as expected.

[0068] The volumetric flow rate Q can now be calculated as:

[00002]$\begin{array}{cc}Q=\frac{\pi D^2}{4}V& \left(2\right)\end{array}$

using the measured diameter D and the measured velocity V as shown above. The results for the three orientations are shown in FIG. 10. The line 24 is again a one-to-one correspondence between the measured and the set values as a guide. These results provide validation of the ability of the Doppler OCT technique used here to precisely quantify flow parameters in a configuration relatively similar to the blood vessels embedded in retinal tissue.

[0069] Line-scanning Doppler flowmetry measurements were performed for all orientations and set velocities as in the OCT measurements simultaneously with the OCT scans. Ten datasets were acquired for each measurement configuration and were averaged assuming constant flow during the measurement time (˜6.26 s). The measurements were also averaged along the tube over a set distance which was selected to avoid strong specular reflections which generally skew the results.

[0070] FIG. 11 shows the results of the Velocity measurements with LSDF for the three orientation angles. One can notice in FIG. 11 that measured Velocity shows a linear increase for a certain set velocity range, and then it saturates, it flattens out. For visualization purposes it is fine, but for quantification purposes it is a problem and it needs to be addressed. Below we present a potential explanation of the phenomenon observed and a solution to velocity quantification using OCT.

[0071] To better understand the issue, we looked at the changes of the power spectrum with increasing flow velocity. FIG. 12A shows the realigned (flattened) volume image of the tube. FIG. 12B shows a few power spectra across the flow. The spectra were averaged along the flow and were normalized to the average spectrum over the entire image. The average spectra 5 and 6 are close to the center of the tube, 4 and 7 are above or below the center, while 1-3 and 8-11 are at the edge or outside the tube. There are 128 sampling points in the time trace of intensity fluctuations and the Fourier transform of the time trace provides 64 points in the power spectrum (the horizontal axis in FIG. 12) as frequencies. The maximum measurable frequency is 45 kHz (half the sampling rate) shown at point 64. The minimum measurable frequency is 0.7 kHz (point 2, point 1 being the DC).

[0072] The velocity is defined in LDF as the mean frequency within the measurement range which works well for slow flow. The problem starts as the flow speed increases and the bell shape starts to move to the right. At some point it starts clipping on the right side as one can see in the sequence shown in FIG. 13 as the set velocity increases. For slow flow, the largest Doppler shift frequencies are still within the measurement window and the mean frequency can be properly evaluate from the power spectrum. This corresponds to the linear increase in FIG. 11 for slow flow. As the flow speed increases, the Doppler shift frequencies increase and the bell shape of the power spectrum moves to the right. At some point, the largest frequency shifts exceed the maximum measurable frequency with the current system of 45 kHz. As the power spectrum is clipped on the right side and a significant part of the power spectrum cannot be measured, the average frequency of the measured spectrum is clearly underestimated and the dependency shown in FIG. 11 flattens out at a value of the order of 25 kHz.

[0073] With this explanation in mind, we developed a solution for proper quantification of Velocity measurement with LSDF. The average of the three orientations is shown in FIG. 14 together with a fit with an error function that seems most appropriate for the data. Since the velocity measurement was validated above with OCT measurements against the set velocity, this fit of LSDF Velocity data as a function of the set velocity can now be used to “correct”, to calibrate the LSDF Velocity measured in kHz into mm/s values. Therefore, there is no single linear calibration factor from Hz to mm/s. We can determine the inverse function of this fit and then extract the correct velocity value [mm/s] for any new measured LSDF Velocity [Hz] value as a lookup table procedure.

[0074] It should be noted here that in the plateau region, which is not perfectly flat, there is still a slow trend up, small errors in the Velocity measurement result in large swings in the calibrated velocity value. It should also be obvious that measured Velocity values larger than the largest value of the fitting curve generate invalid correction and cannot be used.

[0075] Using the inverse function of the fit shown in FIG. 14 we can convert the measured Velocity [kHz] values into calibrated velocity [mm/s] value. The results are shown in FIG. 15.

[0076] Similarly to the estimation of the tube diameter from OCT data, we can evaluate the diameter from LSDF data. The need for that stems from the fact that the OCT measurement for flow analysis is a local measurement, over the small circular scan, while LSDF is providing a large area map of the flow. Lateral average of the Volume image provides a profile of the Volume across the tube as shown in FIG. 16A. Volume is calculated as the total power of intensity fluctuations relative to the DC (non-Doppler shifted) value and is generally regarded as the number of scatterers within the laser interrogation volume. Therefore, Volume is expected to drop to zero at the tube wall. The profile in FIG. 16A is fit with a parabolic profile and the tube diameter is estimated from the distance between the two points on the horizontal axis (scaled in μm) where the fit drops to zero.

[0077] The results of the tube diameter measurement with LSDF for the three orientations are shown in FIG. 16B. As with other measurements, the errors are more significant below 5 mm/s set velocity.

[0078] Having both the velocity (calibrated) and the diameter (average over the three orientations), we can calculate the volumetric flow rate Q using equation 2 and we can compare it with the set flow rate and the OCT measured flow rate. FIG. 17 shows the LSDF flow rate as a function of the set flow rate. The results show a quite remarkable linear dependence.

[0079] FIG. 18 shows the comparison between the volumetric flow rate measured with OCT and LSDF and it clearly validates the concept of calibrating the LSDF measurements using the OCT measurements.

[0080] The ability of the Doppler OCT technique used here to precisely quantify flow parameters in a configuration similar to the blood vessels embedded in retinal tissue has been tested and validated using a glass tube at three different orientations with respect to the OCT beam. OCT is used as a local calibration probe. LSDF is used to generate large area maps of the retinal blood low. LSDF can also be used to measure the diameter and the velocity of the flow and the OCT measurement can be used to calibrate the LSDF measured velocity into the proper unit of measurement [mm/s].

[0081] Following the experiments described above on microfluidics that validated the ability of the combined OCT-LSDF technology to quantify the flow parameters, preliminary demonstration on the eye of human volunteers and on rats with retinal degeneration was performed.

[0082] Ten LSDF large area raster scans were acquired simultaneously with 200 circular OCT scans. The OCT scans were positioned to intersect a retinal blood vessel identified live in the SLO image or in the LSDF Velocity map following the concept shown in FIG. 2. Since the optimum position of the OCT circular scan is such that it intersects the blood vessel along a diameter of the circle, the distance between the two spots should be approximately half the lateral image size. Therefore, the operator adjusts the relative position of the circular scan with respect to the blood vessel watching the position of these two spots after an initial alignment brings the circle close to the blood vessel in the SLO/Velocity image.

[0083] FIG. 19 shows two examples of OCT reflectivity (top)/phase (center) images that were aligned and averaged over the 200 acquired scans. One can clearly identify the two spots corresponding to the location of the blood vessel both in the reflectivity and in the phase image.

[0084] It should be noted here that the left image shows a white and black spot while the right image shows two white spots. White vs. black indicates that the axial component of the flow velocity is directed up or down. The bottom row in FIG. 19 shows an illustration of the scan geometry that explains this effect. The angles in these diagrams are greatly exaggerated for illustration purposes. In a real scan configuration, the total angle at the top of the scanning cone is ˜2°. The left image in FIG. 19 illustrates the situation when the axis of the scanning cone is almost normal to the flow and therefore the axial velocity component (along the OCT beam) can be towards or away from the tip of the cone and can have a positive or a negative Doppler shift (phase shift). The right image in FIG. 19 shows a situation in which both velocity components are towards the tip of the cone having the same sign. One can notice that the black/white spots in the left image have approximately the same height (distance from the scan pivot point) while the two white spots in the right image have different heights, both situation resembling the geometry shown in the diagrams in FIG. 19.

[0085] The example shown in FIG. 19 demonstrates the advantage of the small circular scan that intersects the blood vessel twice at different angles as compared to a single B-scan that intersects the blood vessel only ones. A single B-scan normal to the blood vessel cannot measure Doppler shift. The circular scan used here can never have both intersections normal to the vessel simultaneously; therefore, the circular scan configuration can provide the absolute velocity in any orientation of the flow velocity.

[0086] Processing software automatically identifies the two spots (illustrated in FIG. 20), calculates the height difference and therefore the tilt angle as described above for the glass tube. A parabolic fit is performed on the identified phase spots to measure the diameter of the blood vessel and the axial velocity component. Knowing the tilt angle and the axial velocity component, the absolute velocity is calculated. The circular overlay in the LSDF Velocity image (left in FIG. 20) allows us to estimate the LSDF Velocity at the same location as that of the OCT scan and therefore to calibrate the LSDF Velocity image using the calibration function determined from the microfluidics calibration procedure.

[0087] In one example, the imaging instrument is illustrated through a block diagram in FIG. 21, an optical design diagram in FIG. 22 and CAD diagrams in FIG. 23. The LSDF/OCT imaging system is based on two main optical paths: LSDF and OCT imaging paths. The LSDF imager uses a fixed cylindrical lens and a slit/strip mirror to generate a line of light which is scanned with a single galvanometer scanner (LSDF scanner). The OCT imager uses a pair of galvanometer scanners to steer the OCT beam independently from LSDF. The LSDF and OCT optical paths are then combined at a dichroic beam splitter D1. The common path proceeds to the scan lens group S and to an ophthalmoscopic lens (VOLK lens up to 66 D) which together with the eye optics relays the image plane to the retina.

[0088] The LSDF path begins with the LSDF collimator (located on the back of the plate shown in FIG. 23) which produces a 10 mm beam. The LSDF beam is then passed with a turning mirror to the central section optical plate where a rotary mount holds a 15 mm focal length cylindrical lens which focuses the beam to a line near the aperture-separating strip mirror SM as shown in FIG. 23B. This line is scanned by the LSDF galvo, reflected by the LSDF/OCT combining dichroic D1, and relayed by the scanning/imaging optics (and the eye optics) to the retina.

[0089] The strip mirror SM is a 2 mm width section of a 1 in. plane mirror that reflects the focused LSDF line beam and passes the returning reflection from retinal focal plane over the whole aperture. The LSDF scanner is conjugate to the ocular pupil (approximately 3 to 5 mm from the corneal surface) while the strip mirror is nearly conjugate to the corneal surface. This feature ensures that reflections from cornea are efficiently stopped by the strip mirror leaving as much as 80% of the collection aperture for gathering the reflected light from the retina, and therefore usually requiring no other means of eliminating unwanted reflections (e.g., polarizers).

[0090] The LSDF optical detection path begins from the eye model at right and proceeds through the VOLK ophthalmic lens; through the front scan lens group S (achromat and negative meniscus); reflects from the LSDF/OCT beam-combining dichroic D1 and the LSDF scanner; passes the strip mirror SM to the line-scan focusing lens LS; and reflects from the turning mirror M to the line-scan camera (CCD).

[0091] The OCT imaging path consists of a triplet collimator C (Thorlabs—25 mm focal length and ˜5 mm beam diameter) and a pair of x-y galvo scanners SC (OCT H and V). The collimated beam passes through the LSDF/OCT beam-combining dichroic D1 and then to the retina through the imaging optics common to the LSDF path. The OCT collimator C and the compound lens S define the focal plane to which the imaging path of the LSDF channel needs to be focused during instrument alignment to ensure that the depth range of the LSDF and the OCT channels overlap in the retina. The OCT detection path also includes the OCT interferometer and the spectrometer.

[0092] Although specific features of the invention are shown in some drawings and not in others, this is for convenience only as each feature may be combined with any or all of the other features in accordance with the invention. The words “including”, “comprising”, “having”, and “with” as used herein are to be interpreted broadly and comprehensively and are not limited to any physical interconnection. Moreover, any embodiments disclosed in the subject application are not to be taken as the only possible embodiments.

[0093] In addition, any amendment presented during the prosecution of the patent application for this patent is not a disclaimer of any claim element presented in the application as filed: those skilled in the art cannot reasonably be expected to draft a claim that would literally encompass all possible equivalents, many equivalents will be unforeseeable at the time of the amendment and are beyond a fair interpretation of what is to be surrendered (if anything), the rationale underlying the amendment may bear no more than a tangential relation to many equivalents, and/or there are many other reasons the applicant cannot be expected to describe certain insubstantial substitutes for any claim element amended.

[0094] Other embodiments will occur to those skilled in the art and are within the following claims.