Patent Application: US-201213531034-A

Abstract:
the present invention provides for the detection and display of polarization scrambling tissue without resolving the polarization state of the backscattered imaging beam . in one embodiment , we illuminate the tissue using two different polarizations . a first polarization determines a first image of high intensity while the second polarization determines a second image of low intensity . comparison and combination of the first and second images determines tissue which scrambles the polarization in neighboring detection cells .

Description:
the embodiments , examples and descriptions illustrate the principles of the invention and its practical applications and are not a definition of the invention . modifications and variations of the invention will be apparent to those skilled in the art . the claims define the scope of the invention and include known equivalents and equivalents unforeseeable at the time of filing of this application . one embodiment of the invention is an apparatus for computing a tomographic image of a depolarizing tissue . one such apparatus includes an optical coherence tomography ( oct ) device comprising of an interferometer like the one depicted in fig1 having a source arm 105 , a reference arm 112 , a sample arm 117 ( here shown in two parts since the splitting coupler 110 and combining coupler 130 are separate and distinct ) and a detector arm 135 . a source 101 , typically a superluminescent diode ( sld ), of at least partially spatially coherent light is coupled to the source arm 105 . a polarization controller 140 , capable of varying the polarization state of light within a limited range , is coupled to the sample arm 117 of the interferometer . the sample 120 is scanned with light via scanner 125 and light returned from the sample arm interferes with light from the reference arm in coupler 130 . the interference is detected 150 , forming first intensity data . intensity data may be detected using either time domain techniques or frequency domain techniques . typically , the z - axis is chosen along the beam - line of the optical illumination . data acquired along a beam - line is often referred to as an a - scan . the scanner 125 causes the beam - line to vary transversely . the transverse directions are generally called x and y , though their relative orientation and location depending upon the choice of orientation of the z - axis and the location chosen for the origin . while the interferometer of fig1 is a mach - zehnder transmissive reference arm architecture , the interferometer could also be a michelson architecture , shown in fig2 . the michelson architecture replaces the transmissive reference arm with a reference arm with reference reflector 160 . the oct system containing the interferometer should acquire data rapidly . a frequency domain oct system is preferred . a frequency domain system may be either a spectral domain system , including a wide - band illumination source and a spectrometer , or a swept source system wherein narrowband frequencies are swept across the frequency band . for example , the spectral domain oct system described in u . s . patent application ser . no . 11 / 820 , 773 , filed jun . 20 , 2006 ( publication 2007 / 00291277 ) and incorporated herein by reference can be readily modified to support this invention . as an alternative to varying the polarization in the sample arm 117 , the polarization can be varied in the reference arm 112 by moving the polarization controller from the sample arm to the reference arm . in one embodiment , the tissue is illuminated twice and the polarization controller is set to impart two different polarizations on light passing through it . in this case , as depicted in fig3 , we first establish two polarizations 210 . in one embodiment , we illuminate the tissue using a sequence of polarizations and determine the polarizations which create a maximum and minimum in the total intensity signal . we then acquire first and second images 220 using the polarizations which achieved the maximum and minimum intensities . when scanning can be performed quickly enough , the number of scans performed with uniquely different polarization states can be large , on the order of 25 , 50 or even more different polarizations . however , when scanning time is limited , the number of trial scans with different polarization states may need to be kept quite small . in the latter case , the number of scans at different polarization states can be as small as 4 or 5 . it should be appreciated that when the number of scans is smaller , the variation in polarization state for each scan should likely be larger than if a large number of scans can be accommodated . comparison and combination of the first and second image intensity signals 230 enables detection of the depolarizing tissue . optionally , this image is displayed 240 or stored ( not shown ). in one embodiment , the same tissue is scanned to establish the two polarizations intended to be used in scanning for the final depolarization image . for example , if the depolarization image sought is a b - scan of a region of the eye and the total intensity used to determine the maximum and minimum requires scanning the entire b - scan , then the images acquired in 220 are optimally saved during the procedure 210 used to establish the scanning polarizations . this is readily accomplished when , for each polarization used in procedure 210 , the b - scan is acquired , its total intensity is computed and compared to the maximum and minimum previous total intensities of previous b - scans . if it is greater than the previous maximum , its intensity value becomes the new maximum intensity value and the image replaces the previous maximum intensity image . if it is less than the previous minimum , its intensity value becomes the new minimum intensity value and the image replaces the previous minimum intensity image . alternatively , if the polarizations are established by scanning over a limited region , procedure 210 may quickly test a number of polarizations , choose two and then acquire images 200 in a completely separate procedure . in one embodiment , the polarization is varied over a sampled subset of all polarizations attainable by a single polarization paddle . the polarization paddle may be located in the sample arm or in the reference arm . as the polarization paddle is rotated , it imparts different polarizations on the light traveling through the fiber mounted on the paddle . in the representation of fig4 , a fiber would be mounted with a u - shape bend onto the paddle , but it could otherwise form the more traditional circular loop . the paddle can rotate out of the plane of the paper with the u - shape remaining in one plane at all times . the paddle design parameters are the three radii ( r 1 , r 2 and r 3 ) and the three angles ( α 1 , α 2 , α 3 ). these parameters are typically chosen to compensate for system birefringence . in another embodiment , the polarization is varied over a sampled subset of all polarizations attainable by two or more polarization paddles . alternatively , the polarization may be varied using a liquid crystal based polarization controller or an electro - optical polarization controller . other polarization controllers with substantially similar operating parameters may be used , as will be clear to those versed in the art . in all cases , the polarizations are varied to determine the polarizations that establish detection efficiencies for the two images . in another embodiment , the establish scanning polarization process 210 of fig3 is fixed in hardware . this hardware is a fast polarization modulator designed to rapidly vary the polarization to sample the scattering with a fixed number , say n , of detection efficiencies . this hardware may be placed in either the sample arm or the reference arm of the oct interferometer . for example , this hardware may be a rotating retarder of n positions , for example a traditional ¼ lambda or ½ lambda rotating waveplate . any polarization modulator that can produce a finite number of diversely varying polarization states where the polarization states are essentially fixed for each a - line may be used . typically , n is less than 20 and preferentially it is between 2 and 6 . preferably the polarization is fixed or nearly fixed during acquisition of each a - line . successive a - lines have different polarizations applied in the sample arm . while a - lines are normally quite closely spaced in typical oct imaging , the a - lines can be oversampled or the same tissue may be imaged n times using a different polarization setting for each a - line acquisition . preferentially , the a - lines are closely spaced . the acquired data 220 may be processed as n interleaved images , each acquired with a different polarization setting . from these n images , we select the maximum intensity image and the minimum intensity image from which to compute the depolarization image 230 . the system hardware may alternatively be placed in the reference path . alternatively , a polarization paddle or other polarization controller may be placed in the reference arm . advantageously , this paddle or controller may be varied to increase the variation between the maximum intensity image and the minimum intensity image . preferably , this optimization is performed over a small region and then the polarization paddle is set for full image acquisition . this decreases the full image acquisition duration . however , when the total time of acquisition is not critical , the polarization paddle may be varied over the full image region to optimize its setting . the polarization paddle and the n polarization state fixed hardware may be in the same interferometer arm or in the alternate arms . in yet another embodiment , a first polarization is a priori selected to be at or near the maximum for a population and a second polarization is a priori selected to be at or near the minimum for the same population . the a priori selection may be heuristically determined from a sample set , derived analytically from a model , or obtained by other means . in one embodiment , the maximum intensity signal is the maximum intensity averaged over an entire b - scan . in this embodiment , the minimum intensity signal is the minimum intensity averaged over an entire b - scan . alternatively , the maximum signal intensity may be determined by the average intensity over a region associated with a particular structural feature in the eye such as near the inner limiting membrane ( ilm ) within a b - scan . in this case , the b - scan is acquired for a particular polarization , the b - scan is segmented to locate the ilm , and the region near the ilm is identified before the signal intensity is computed . in this embodiment , the minimum intensity signal is also determined over the same region near the inner limiting membrane ( ilm ) within a b - scan by computing the average intensity over the region . in this case , the b - scan is acquired for a particular polarization , the b - scan is segmented to locate the ilm , and the region near the ilm is identified before the signal intensity is computed . in order to reduce the computation time , the b - scan may contain only a limited number of a - scans . indeed , either method can be implemented using only a single , representative a - scan instead of the full b - scan . if the maximum signal intensity is the maximum average intensity of a b - scan over all polarizations , then the minimum signal intensity is the minimum average intensity of a b - scan over all polarizations . similarly , if the maximum signal intensity is the maximum average intensity near the ilm over all polarizations , then the minimum signal intensity is the minimum average intensity near the ilm over all polarizations . that is , the minimum signal intensity should be computed over the same or over nearly the same polarizations as the maximum and computed over the same or nearly the same regions of tissue as the maximum . preferentially , the images are scanned interleaved , to reduce or eliminate motion artifacts . however , it may be impractical to interleave the images . in this case , proper registration of a - lines between images reduces motion artifacts that distort the final image . in one embodiment , the images are combined on a pixel - by - pixel basis . let i mn + represents the intensity of the ( m , n ) pixel of the image acquired using the polarization associated with the maximum intensity and i mn − represents the intensity of the ( m , n ) pixel of the image acquired using the polarization associated with the minimum intensity . if i mn represents the total intensity and i mn p represents the polarized intensity then the degree of polarization , mn = i mn p / i mn . since , for perfectly correct polarization ( i . e ., where i mn + contains all of the polarized intensity and ½ of the unpolarized intensity ), and , for perfectly orthogonal polarization ( i . e . where i mn − contains ½ of the unpolarized intensity ) dod = 1 - mn = η mn = 2 ⁢ i mn - i mn + + i mn - . mn is a measure approaching 1 where the first and second image intensities are nearly equal and approaching 0 where the first image intensity is much larger than the second image intensity . if a single pixel is large enough that sufficiently many scatters are within an imaging cell ( pixel ), and then η mn represents the degree of depolarization for each pixel of an image . however , if each detection cell is sufficiently small that it represents a single scatterer or the imaging technique is coherently detected , then η mn should be computed using a smoothed i mn + and i mn − , where smoothing is performed over a window sufficiently large to account for the number of scatterers needed to depolarize the incoming light . in particular , for oct , where the detection is confocal and coherent and the detection cell is approximately the size of speckle , the image should be smoothed over a region sufficiently large to cover a statistically meaningful number of different speckle cells . that is , i mn + = ∑  j - m  ≤ j  k - n  ≤ k ⁢ ⁢ ( w ⁡ ( j - m , k - n ) ⁢ i ^ mn + ) and i mn - = ∑  j - m  ≤ j  k - n  ≤ k ⁢ ⁢ ( w ⁡ ( j - m , k - n ) ⁢ i ^ mn - ) , where w is a smoothing weight , î mn + and î mn − are the measured intensities , and j and k govern the size of the window . the weight w = 1 simply pixel - wise averages the intensities . windows with even length boundaries are also anticipated and readily understood by those versed in the art . the dod η mn can be used create a color image by modulating the hue , saturation , or value of an hsv color representation of the image . for example , the hue , saturation , and value may be set as a functions of η mn , i mn + , and / or i mn − . in one instance , the hue may be set to a function of η mn while the saturation is saturated and the value is set to a function of i mn + . fig5 shows an example of a color image of depolarizing tissue . fig5 a is an image of tissue imaged with a polarization paddle set to achieve a near maximum intensity i mn + . in order to view the dynamic range , the image is essentially log ( i mn + ). fig5 a is displayed in reverse video , with high intensity shown in black and low intensity shown in white . fig5 b , also shown in reverse video , is the same tissue imaged with a polarization paddle set to achieve a near minimum intensity i mn − . fig5 c is a color representation of the depolarization image where the hue is set to magenta , the saturation is set to the degree of depolarization and the value is set to the logarithm of i mn + . alternatively , the hue may be set to a function of η mn while the saturation is saturated and the value may be set to a function of i mn − . alternatively , the intensity the image i mn can be used create a grayscale image . fig6 shows an example of a grayscale image of depolarizing tissue . fig6 a is the same image as fig5 a . it shows tissue imaged with a polarization paddle set to achieve a near maximum intensity i mn + in reverse video . fig6 b , also shown in reverse video , is the same as fig5 b . this is the same tissue imaged with a polarization paddle set to achieve a near minimum intensity i mn − . fig6 c is a grayscale modulation of the degree of depolarization with the image i mn − . a grayscale modulation may be computed as : ĩ mn + = f ( η mn ) g ( i mn + ), or ĩ mn − = f ( η mn ) g ( i mn − ), or more generally î mn = h ( i mn − , i mn + ). since η mn is likely very nearly 1 in regions of weak signal , η mn by itself enhances some noise . modulation of η mn should maintain its strength where i mn − is near its maximum , while reducing its strength where i mn − is near its minimum . the normal image display for oct is essentially logarithmic ( in order to increase the dynamic range distinguishable by the human eye ). fig6 c is an exemplary embodiment of the modulation η mn log ( i mn − ). in general , while f and g may well be the identity function , it is preferred that g vary more slowly through values where i mn − is rich in signal , such as a logarithmic function or even a step function , thresholded at a known noise level . a continuous function transitioning rapidly from nearly 1 above a threshold to nearly 0 below the threshold provides an alternative to a true step function . various other modulations of color and grayscale representations are possible and will be appreciated by one versed in the art . a tomographic image composed of a - line scans of enhanced regions of depolarizing tissue can be formed . in some instances , the metric ĩ mn − = f ( η mn ) g ( i mn − ) used to create a depolarizing tissue image is sufficiently dominated by g ( i mn − ) that the image i mn + is unnecessary for computing an approximate depolarizing tissue image î mn − = g ( i mn − ). this is particularly useful when acquiring a 3 - d volume of image data since , once the polarization is determined for imaging i mn − , whether this is done over a b - scan , a portion of a b - scan , or even over an a - scan , the entire 3 - d volume can be acquired using only that fixed polarization . thus , the depolarizing tissue image can be acquired rapidly and without scanning using a distinct second polarization . when only two polarizations are used , it is obvious that there is only one scanning sequence : the sequence that is used to provide information about the depolarization of tissue . the preferred polarizations are already chosen . however , when the preferred polarizations need to be determined , there is a need for a scanning sequence to generate information needed to choose preferred polarizations ( e . g . the polarizations which produce the maximum and minimum average intensity image information ). this scanning sequence used for determining preferred polarizations need not be the same as the scanning sequence used to provide information about the depolarization of tissue . for example , the scanning sequence used to determine the best two polarizations might be the scanning sequence used to generate the lower resolution b - scans of u . s . patent publication 2007 / 0216909 while the scanning sequence used to generate the images from the chosen polarizations might be the high resolution scanning sequence of that patent publication . in general , any appropriate sub - region of the region scanned by the scanning sequence used to provide information about the depolarization of tissue may be scanned to determine the preferred polarizations . in fact , even regions near the target region may be used to choose preferred polarizations , so long as the tissue is sufficiently uniform that the estimate obtained from the polarization choosing scans is relevant to the region scanned to provide information about the depolarization of the tissue . it should be understood that the embodiments , examples and descriptions have been chosen and described in order to illustrate the principles of the invention and its practical applications and not as a definition of the invention . modifications and variations of the invention will be apparent to those skilled in the art . the scope of the invention is defined by the claims , which includes known equivalents and unforeseeable equivalents at the time of filing of this application . 2007 / 0216909 everett et al ., methods for mapping tissue with optical coherence tomography data 2007 / 0146632 chipman , advanced polarization imaging method , apparatus , and computer program product for retinal imaging , liquid crystal testing , active remote sensing , and other applications . u . s . pat . no . 7 , 286 , 227 zhou , et al . method and system for removing the effects of corneal birefringence from a polarimetric image of the retina u . s . pat . no . 7 , 016 , 048 chen et al . phase - 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