the present subject matter provides an improvement on 3d optical angiography . single - pass volumetric bidirectional blood flow imaging ( spfi ) sdoct detects moving scatterers using a modified hilbert transform without the use of spatial frequency modulation . since no frequency modulations are required and spfi processing is applied to the spatial frequency content across a single b - scan , spfi is applicable for both spectrometer - based and swept - source oct systems , provided they have comparable b - scan acquisition rates . unlike previously described techniques , which require two separate b - scans to detect positive and negative flow , each with modulations tuned to the desired flow direction , spfi is able to resolve bidirectional flow in a single b - scan pass across the sample . the depth - encoded complex spectral interferometric signal from m discrete sample reflectors for an sdoct system can be written as follows : s ⁡ [ k , x ] = ∑ m = 1 m ⁢ a m ⁢ exp ⁡ [ ⅈ ⁡ ( 2 ⁢ nk ⁢ ⁢ δ ⁢ ⁢ z m + θ h ⁡ [ x ] + θ m ⁡ [ x ] ) ] ( 1 ) where k is the wavenumber and a m , δz m , and n represent the reflectivity , depth , and refractive index of each reflector , respectively . further , θ m is a spatially - varying phase term related to the doppler frequency - shift of the axial component of scatterer motion measured relative to the first lateral a - scan across each scatterer . θ h [ x ] represents the sum of phases arising from an optically heterogeneous sample and can be represented as a random variable with a mean of zero and a standard deviation ( ρ h [ x ]) that is related to the sampling spot - size . previous work has demonstrated in vivo 3d blood flow mapping by imposing a known axial phase component across a lateral scanning dimension ( b - scan ) using a moving reference arm . similarly , modulation of this phase component has also been used for full range complex conjugate resolved sdoct reconstruction . spfi - sdoct is a modification of these previous techniques where scatterers moving above a threshold velocity are imaged without the use of a modulation frequency . the recorded interferometric signal represents the real - part of a summation of signals from m discrete reflectors ( eq . ( 1 )) in a coherence volume ( i . e . spot - size × coherence - length ) and thus the phase term ( θ m [ x ]) can be represented as follows : θ m ⁡ [ x ] = 2 ⁢ nk ⁢ ⁢ π wo / 2 ⁡ [ x - dl d ] ⁢ v m v x ⁢ ( x - dl d + w o 2 ) ( 2 ) where v m and v x are the axial components of scatterer velocities and lateral scan speed , respectively . π is the boxcar function defined by π wo / 2 [ x − dl / d ]= h [ x − dl / d + w o / 2 ]− h [ x − dl / d − w o / 2 ] and restricts the measured phase to moving scatterers within the spot - size , w o . d represents the a - scan location of the moving scatterer across a lateral scan of length l , sampled with a density of d a - scans . combining eq . ( 1 ) and eq . ( 2 ), the sampled interferometric signal can be represented as the conventional sdoct signal with a velocity - associated phase modulation : s ⁡ [ k , x ] = ⁢ ∑ m = 1 m ⁢ a m ⁢ exp ⁡ [ ⅈ ⁢ ⁢ 2 ⁢ nk ⁢ ⁢ δ ⁢ ⁢ z m ] ⁢ exp ⁡ [ ⅈ ⁢ ⁢ θ h ⁡ [ x ] ] ⁢ exp ⁢ [ 2 ⁢ ⁢ ⅈ ⁢ ⁢ nk ⁢ ⁢ π ⁢ ⁢ w o / 2 ⁡ [ x - dl d ] ⁢ v m v x ⁢ ( x - dl d + w o 2 ) ] = ⁢ ∑ m = 1 m ⁢ a m ⁢ exp ⁡ [ ⅈ ⁢ ⁢ 2 ⁢ ⁢ nk ⁢ ⁢ δ ⁢ ⁢ z m ] ⁢ g ⁡ [ x ] ⁢ h ⁡ [ x ] ( 3 ) where g [ x ] and h [ x ] are the interferometric components associated with the optical heterogeneity and moving scatterers , respectively . the recorded spectral datacube ( l [ k , x , y ]) in sdoct is comprised of the real - parts of the interferometric signals ( eq . ( 3 )) accumulated during a 2d raster scan and can be separated into a series of b - scan ( l [ k , x ]) data slices . for instance , fig1 ( a ) illustrates a spectral inverse fourier transform of a b - scan . the inverse fourier transform in the spectral dimension , ft − 1 k → z [ i [ k , x ]]= r [ z , x ]+ r [ z , x ], represents the depth - resolved reflectivity map in conventional sdoct . an example of a conventional sdoct depth - resolved reflectivity map and lateral fourier transform yields are shown in fig1 ( b ) and 1 ( c ). fourier transforming the lateral dimension ( i . e ., b - scan ) yields the spatial frequency content : ft x -& gt ; u ⁡ [ i ⁡ [ k , x ] ] = ∑ m = 1 m ⁢ a m ⁢ cos ⁡ [ 2 ⁢ nk ⁢ ⁢ δ ⁢ ⁢ z m ] ⊗ re [ g [ u ] ] ⊗ re [ h [ u ] ] ( 4 ) h [ u ] is the spatial frequency content of the phase associated with moving scatterers integrated across a spot size re [ h ⁡ [ u ] ] = ⁢ re [ ∫ d ⁢ ⁢ l / d - ⁢ w 0 / 2 d ⁢ ⁢ l / d + ⁢ w 0 / 2 ⁢ exp [ 2 ⁢ ⁢ ⅈ ⁢ ⁢ nk ⁢ ⁢ v m v x ] ( x - dl d + w o 2 ) ⁢ exp ⁡ [ - ⅈ ⁢ ⁢ ux ] ⁢ ⅆ x ] = ⁢ κ ⁢ cos ⁡ [ w 0 ⁡ ( u - 2 ⁢ nk ⁢ ⁢ v m v x ) ] - 1 u - 2 ⁢ nk ⁢ ⁢ v m v x ( 5 ) where κ is a constant scaling factor . eq . ( 4 ) can be represented in the form of ft x → z [ i [ k , x ]]= r ′[ u , k ]+ r ′[ u , k ]+ v ± [ k , u + f d ,± ]. here both moving ( v ± [ k , u ÷ f d ,± ]) and non - moving ( r ′[ u , k ]) scatterers in the sample are imaged ( see , e . g ., sample spatial frequency information in fig1 ( d ) and 1 ( e )), where f d ,± represents the doppler frequency shift associated with the axial components of positive and negative scatterer motion . note that all scatterer motion and associated flow refer to the axial components of motion ( e . g . negative flow refers to the axial component of velocity for scatterers moving away from the sample beam ). the doppler frequency shift is related to eqs . ( 2 ) and ( 3 ) by f d = nkv m / π and can be considered a sum of frequency shifts , representing all moving scatterers within a coherence volume , convolved with their respective reflectivities , in the spatial fourier domain . non - moving scatterers , in this case , are centered around dc ( see , e . g ., fig1 ( d ) and 1 ( e )). the spatial frequency bandwidths of both moving and non - moving scatterers across a b - scan are related to the spatial frequency standard deviation of the heterogeneity term ( g [ u ]), where the standard deviation results from a summation of independent random variables : σ h ⁡ [ u ] = ∑ x = 1 l ⁢ σ h 2 ⁡ [ x ] ( 6 ) and is representative of the spatial correlation between sequentially sampled a - scans with a high correlation lower limit associated with the optical heterogeneity of scatterers within a coherence volume ( θ h [ x ]) and a nyquist sampling upper limit . since the recorded spectral datacube is real - valued , the previously described spatial fourier transforms include both real and complex conjugate peaks . previous works have shown that by imposing a constant phase shift between a - scans along a single lateral scan direction , a carrier frequency can be added to the phase term and thus isolating real and complex conjugate peaks to opposite spatial frequency spaces . windowing out the conjugate peaks , inverse transforming back to k - space ( e . g ., using a hilbert transform ), and then applying conventional spectral fourier transforms on the data allowed for separation of moving ( e . g ., flow ) and non - moving ( e . g ., structure ) scatterers to opposite image half - planes . imaging of bidirectional flow required applying positive and negative carrier frequencies and processing each datacube separately . in spfi - sdoct , we recognize that without the use of carrier frequencies , the spatial frequencies of moving and non - moving scatterers do not overlap at spatial frequencies above the non - moving scatterer bandwidth . for instance , fig2 ( a ) and 2 ( b ) illustrate that a lateral fourier transform of a b - scan yields a spatial frequency of a sample centered around dc and a spatial frequency of moving scatterers shifted by their respective doppler frequencies . an analytic signal for the spectral interferogram can be obtained by applying a heaviside function ( h [ u − f t ]), frequency - shifted outside of the structural bandwidth ( see , e . g ., fig2 ( c )), and then inverse fourier transforming the result ( see , e . g ., fig2 ( d )). application of this modified hilbert transform ( ht *) enables bidirectional flow imaging by windowing eq . ( 4 ) to yield the following : ft x -& gt ; u ⁡ [ i ⁡ [ k , x ] ] ⁢ ⟶ h ⁡ [ u - f t ] = α ⁡ ( v + ⁡ [ k , u + f d , + ] + v _ - ⁡ [ k , u - f d , - ] ) ( 7 ) where α represents the fractional portion of bidirectional flow with doppler frequencies outside of the spatial frequency bandwidth of non - moving scatterers ( see , e . g ., fig2 ( c )). this threshold frequency ( f t ) defines the minimum detectable velocity in spfi - sdoct and is related to the spatial correlation of sequential a - scans . therefore , lateral oversampling of a - scans allows for a reduced threshold frequency and thus , slower minimum detectable velocities . all spatial frequencies above f t can be detected given that their associated accumulated phases are above the system phase - noise floor . spatial oversampling and velocity resolution can be related by combing the velocity - related doppler frequency shift with the spatial frequency resolution , which can be represented as follows : dv = l ⁢ ⁢ λ 0 2 ⁢ nd ⁢ ⁢ τ ⁢ ⁢ w o ⁢ cos ⁢ ⁢ θ d ( 8 ) where λ 0 is the center wavelength , τ is the integration time , and θ d is the doppler angle between the scanning beam and the direction of scatterer motion . further , n is the index of refraction , l is the lateral scan length , d is the number of a - scans acquired across the lateral scan , and w o is the scanning beam spot size . eq . ( 8 ) shows that velocity resolution in spfi increases and the maximum detectable velocity decreases linearly with increased spatial oversampling . since the resulting complex interferometric signal is a sum of positive moving scatterers and the conjugate of negative moving scatterers , application of conventional sdoct processing yields a flow image where bidirectional flow is imaged to opposite sides of dc . for instance , fig2 ( e ) shows a spectral inverse fourier transform of the analytic interferometric signal maps depth - solved reflectivities of bidirectionally - moving scatterers on opposite image half - planes which can then be overlaid for vessel identification ( see , e . g ., fig2 ( f )). finally , application of the heaviside function to isolate moving scatterers also acts to reduce the overall noise of the velocity map . it is noted that in a similar analysis for complex conjugate removal , it can be demonstrated that spatial frequency windowing provides an snr gain related to the window function . similarly in spfi , the resulting vessel map should benefit from a signal gain as a result of rejection of noise components outside of the velocity detection band . in addition , since velocity depends linearly on frequency , application of a spatial frequency window will necessarily filter out all velocities not traveling at velocities described by the following : δ ⁢ ⁢ v = l ⁢ ⁢ λ 0 ⁢ w ⁡ [ u ] nd ⁢ ⁢ τ ⁢ ⁢ cos ⁢ ⁢ θ d ( 9 ) here , w [ u ] represents the spatial frequency fwhm bandwidth of the gaussian window . there is , however , a tradeoff between the spatial frequency window width and the resulting spatial resolution of the velocity - resolved intensity image . for a gaussian window in frequency - space , the resulting spatially resolved vessel map will be convolved with a gaussian in the b - scan dimension with fwhm bandwidth w [ x ]= 4 ln2 /( πw [ u ]). thus , to avoid loss of lateral resolution in the velocity - resolved vessel maps , the spatial frequency window bandwidth ( w [ u ]) needs to be constrained such that the associated lateral blurring function ( w [ x ]) does not exceed the scanning beam spot size , w o . the velocity window limit can be represented as the bandwidth w [ u ]≧ 4 in 2 /( πw o ). given the discrete sampling parameters of spfi - sdoct , the minimum w [ u ] before loss of resolution is a factor of 8 ln 2 / π greater than the spatial frequency sampling rate . combining this result with eq . ( 8 ) gives the minimum resolvable velocity resolution as follows : finally , a datacube ( l [ x , z , v ]) for each b - scan can be formed by shifting the spatial frequency window across the bandwidth of moving scatterers , inverse transforming the result back to functions of l [ k , x ], and then performing a spectral inverse fourier transform . for instance , fig3 ( a ) and 3 ( b ) illustrate that a lateral fourier transform of raw sdoct spectral interferogram b - scan ( see fig3 ( a )) yields a spatial frequency of stationary scatterers centered around dc , and spatial frequency of moving scatterers shifted by their respective doppler frequencies ( see fig3 ( b )). applying a frequency - shifted heaviside step function , spatial frequency windowing , and inverse fourier transforming each frequency range ( see fig3 ( c )) recreates the analytic interferometric signal ( see fig3 ( d )). a spectral inverse fourier transform of the analytic interferometric signal maps depth - solved reflectivities of moving scatterers for each corresponding velocity range into a datacube for each b - scan ( see fig3 ( e )). bidirectional flow is mapped onto opposite image half - planes . summing the datacube across all velocity ranges creates velocity - and depth - resolved b - scans . each v - slice represents a velocity range given by eq . ( 9 ), and the datacube can be summed across the v - dimension to create a single velocity - resolved b - scan . similar to conventional doct , velocity wrapping occurs as the spatial frequency content of moving scatterers wraps across the nyquist sampling upper limit . in this case , the scatterers are mapped to the opposite image half - plane and are therefore represented as moving in the opposite direction . phase unwrapping techniques , similar to those used for doct , can be applied for a singly wrapped velocity profile . higher spatial frequencies that wrap across nyquist and have components greater than the threshold frequency ( f t ) will necessarily lose their velocity components due to spfi windowing . this fundamentally limits the resolvable spatial frequency range to twice the nyquist frequency . by way of specific example , spfi - sdoct can implemented on a high - speed sdoct microscope having a specific central wavelength ( e . g ., at about 859 nm ) and a fwhm bandwidth ( e . g ., about 99 nm ), an arrangement for which is shown in fig4 . the sample arm can be a microscope ( e . g ., a custom - built f / 8 . 5 microscope ) equipped with scanning galvanometers and imaging optics optimized for a given spot - size ( e . g ., about 9 μm ). spfi data can be created using both conventional and common path sdoct configurations by blocking the reference arm of a typical sdoct interferometer to implement a self - referenced imaging scheme to reduce phase - noise such that f t is dominated by lateral sampling . phase - noise between sequential a - scans in a homogeneous phantom can be measured in both conventional and common path configurations . since the lower limit of the resolving power in spfi is proportional to the phase - noise , the common path configuration can be used for animal model imaging . interferometric signals can be captured using a line - scan camera ( e . g ., a 2048 pixel line - scan camera from e2v , ltd .). real - time data acquisition , processing , archiving , and display can be performed , for instance by using a software product ( e . g ., a custom software program from bioptigen , inc .). for example , using a 1 . 3 mw sample beam , the snr measured near dc can be 108 db with an axial resolution of 3 . 29 μm in tissue and a 6 db falloff at 0 . 8 mm . dc removal , k - space resampling , and flow imaging using the modified hilbert transform algorithm can further be computed during post - processing ( e . g ., using matlab from mathworks , inc .). in addition , vessel and structure can be visualized ( e . g ., using amira from visage imaging , inc .). in another aspect , the disclosed methods can be specifically applied to retinal microvessel imaging . velocity - resolved spfi - sdoct can be implemented on a high - speed sdoct retinal imaging system employing a light source with a given central wavelength ( e . g ., about 859 nm ) and a given fwhm bandwidth ( e . g ., about 99 nm ), an arrangement for which is illustrated in fig5 . the sample arm can be a modified slit lamp equipped with scanning galvanometers and relay optics for retinal imaging of subjects . the retinal scanner optics can be designed for a particular transverse resolution ( e . g ., about 15 - 20 μm ), as limited by the optics of the eye , across a field ( e . g ., about 12 × 12 mm ). the reference arm can be dispersion compensated using a water cell and matched optics , and interferometric signals can be captured using a line - scan camera ( e . g ., a 2048 pixel line - scan camera from e2v , ltd .). data acquisition , archiving , and real - time processing and display of image magnitude can be performed , such as by a custom software program . in particular , using a 700 μw sample beam , the snr measured near dc can be 110 db with an axial resolution of 4 . 72 μm in tissue and a 6 db falloff at 0 . 8 mm . dc removal , k - space resampling , and flow imaging using the modified hilbert transform algorithm can be computed during post - processing ( e . g ., by using matlab ). vessels and structure can be visualized ( e . g ., using amira and osa isp from kitware , inc .). further , velocity - resolved bidirectional flow imaging can be validated and compared with conventional doct on a flow phantom ( see , e . g ., fig6 ( a )). in particular , two glass micro - capillary tubes ( e . g ., 1 . 5 mm outer diameter , 0 . 6 mm inner diameter ) can be connected using silastic tubing to a syringe pump and pumped with 1 % liposyn at one or more of 10 μl / min , 20 μl / min , 30 μl / min , and 40 μl / min . the micro - capillaries can then be positioned adjacent to each other on an angled stage such that fluid in the tubes flows in opposite directions in a b - scan cross - section , as is shown in fig7 ( a ) through 7 ( c ), simulating bidirectional flow . a standard minimum b - scan size of 1000 a - scans / frame for a 3 mm lateral scan at an integration time of 50 μs is assumed . faster lateral scan velocities can result in galvanometer jitter , and therefore poor image quality and phase stability . spatial oversampling for spfi is defined based on these scan parameters in the following . a 3 mm b - scan can be acquired with 1800 a - scans / frame with an a - scan integration period of 50 μs , a factor of 1 . 8 increase in lateral oversampling . conventional sdoct processing ( see , e . g ., fig7 ( a )) can show depth - ranged reflectivity of scatterers including structure , flow , and their complex conjugate mirror images . the directionality of flow in each tube is not readily discernable using conventional processing steps ( fig7 ( a ), arrows 71 and 72 ). after applying the modified hilbert transform algorithm , all non - moving scatterer reflectivities ( i . e ., structure ) and mirror images of flow can be resolved leaving only positive and negative flow on opposite image half - planes ( see , e . g ., fig7 ( b )). the phantom structural heterogeneity can be band - limited as a result of lateral oversampling , allowing the frequency - shifted heaviside function to window out only moving scatterers ( eq . ( 6 )). similarly , the mirror images of moving scatterers can also be eliminated by application of the modified hilbert transform , leaving only the real - valued positive flow and complex conjugate negative flow , which can be imaged to opposite sides of dc ( see , e . g ., fig7 ( b )). using the spfi processed image , positive and negative flow are separated and overlaid onto the structural image for visualization ( see , e . g ., fig7 ( c )). given the oversampling parameters and the threshold frequency determined experimentally from the phantom data , the magnitude of the detectable positive and negative flow velocities can be about 0 . 39 - 1 . 12 mm / s . in another exemplary arrangement , b - scans of the phantom can be acquired across a 2 mm scan range with 2500 a - scans / frame for spfi and 1000 a - scans / frame with 4 sequential a - scans at each lateral position for doct . the spfi dataset can be laterally oversampled compared to doct because velocity resolution increases as a function of spot - size overlap on the sample ( eq . ( 8 )). both datasets can be acquired with an integration time of 50 μs . at these sampling parameters , the total imaging time for spfi can be a factor of 1 . 6 times faster than that of doct . spfi parameters can be chosen to demonstrate velocity resolution at the lower limit of the detection range for a given integration time as a comparison with doct . since the velocity resolution of doct is limited by the phase noise of the system instead of spatial sampling parameters , as is the case with spfi , an appropriate lateral spacing can be used to minimize scanner jitter . the number of sequential a - scans used can be indicative of common doct sampling parameters . in addition , a threshold frequency can be determined , which can filter out all spatial frequencies of non - moving scatterers . a gaussian window can then be moved across the remaining spatial frequencies to velocity - resolve the b - scan for flow rate measurements . the shifted window can be set such that the velocity - range resolved has a fwhm and shifts at increments of 24 . 3 μm / s . given the oversampling parameters and threshold frequency window used , the magnitude of the total detectable velocity range for the axial components of positive and negative velocities can be about 0 . 61 - 11 . 53 mm / s . velocity - resolved scatterer information can then be overlaid onto structural b - scans for visualization ( see , e . g ., fig8 ( a )). doct volumes can be processed using standard phase - difference methods . velocity profiles for both doct ( see , e . g ., fig8 ( b ) and 8 ( c )) and spfi ( see , e . g ., fig8 ( d ) and 8 ( e )) can be fit to laminar flow curves , and measured capillary cross - section and flow rate can be calculated for both imaging methods and compared ( see , e . g ., fig8 ( f ) and 8 ( g )). vessel imaging can also be demonstrated on chicken embryo and mouse tumor window chamber models ( see , e . g ., fig6 ( b ) and 6 ( c )). in particular , fertilized hubert ross chicken eggs can be incubated at 38 ° c . and 97 % humidity in a forced - draft incubation chamber . a 3 mm × 3 mm volume mosaic was created by acquiring nine 1 mm × 1 mm volumes imaged with 1800 a - scans / frame and 100 frames / volume , a factor of 5 . 4 increase in lateral oversampling . at hamburger - hamilton ( hh )- stages 23 - 25 , a window can be created through the outer shell and the chorionic membrane can be removed . peripheral yolk vasculature ( see , e . g ., fig6 ( b )) can then be imaged using the commonpath configuration with the top amnion surface as the reference reflector . vessels away from the embryo can be imaged to avoid pulsatile flow as a result of heart beat . embryo temperatures can be maintained using a heat - lamp during the course of imaging . each spfi - sdoct reconstructed frame can be separated into two halves and combined to create bidirectional flow maps with intensities corresponding to the reflectivity of scatterers moving into or out of the a - scan axis . the peripheral yolk vascular network , shown in fig9 , is clearly visible in the chicken embryo model . spfi - sdoct imaged vessels are tubular , as expected , and appear to be confined to a 500 μm layer ( see , e . g ., fig1 ). the large and partially imaged vessel across the volume mosaic ( see fig9 , arrow 91 ) is a shadow artifact arising from imaging through a large vessel in the amionic ( i . e ., the reference reflector ) layer . the well defined portion of this vessel ( see fig9 , arrow 92 ) represents areas where the vessel branches from the yolk surface up towards the surface of the amnion . the faded regions ( see fig9 , arrow 91 ) are areas where the vessel is on top of the amion surface and out of the imaging depth range . yolk surface vessels follow the curvature of the yolk sac and are shown gradually descending away from the reference surface ( see fig1 ). detected vessel sizes ranged from 40 μm ( fig9 , arrow 93 ) to 270 μm ( fig9 , arrow 94 ). “ hazy ” sections ( fig9 , arrow 95 ) are indicative of sample bulk motion which resulted in non - moving structural scatterers being resolved along with flow . chicken embryo volumes were acquired using 1 ms a - scan integration time in order to detect small vessels with minimal flow velocities . the magnitude of the detectable positive and negative flow velocities for the embryo model was 65 . 6 - 168 . 6 μm / s . mouse tumor models can be prepared by surgically implanting a titanium window chamber on the back of anthymic ( nu / nu ) nude mice under anesthesia ( ketamine 100 mg / kg and xylazine 10 mg / kg intraperitoneal ). 4t1 metastatic mouse mammary adenocarcinoma cells can be used . during window implantation , 10 μl of a cell suspension of 5 103 cells can be injected into the dorsal skin flap and covered with a 12 mm diameter # 2 round glass coverslip over the exposed skin ( see , e . g ., fig6 ( c )). animals can be housed in an environmental chamber with free access to food and water and standard 12 hr light and dark cycles . mice tumors were imaged two weeks after implantation using the window chamber surface as a reference reflector . mice can be imaged using the commonpath configuration , self - referenced using the surface of the window chamber as the reference reflector . tumor regions were indentified prior to imaging ( see , e . g ., fig1 , arrow 111 ) and show highly tortuous vessels indicative of neoplastic angiogenesis . surrounding vasculature indicate normal skin fold vessels . detected vessel sizes ranged from 20 μm ( see , e . g ., fig1 , arrow 112 ), which approaches the sampling limit of the microscope , to 110 μm ( see , e . g ., fig1 , arrow 113 ). the volumes were acquired using 2 ms a - scan integration time in order to detect small vessels however these volumes were more sensitive to bulk motion artifacts ( see fig8 , arrow 114 ). the magnitude of the detectable positive and negative flow velocities for the mouse tumor model was 32 . 8 - 84 . 3 μm / s . small noise signals throughout the volumes indicate areas of reference reflector saturation due to small optical reflectivity heterogeneities across the reference window chamber surface . in vivo microvessel imaging can further be demonstrated in a normal human retina . first , an oct volume of the subject ( e . g ., a 10 × 10 mm volume ) can be acquired , allowing for the reconstruction of a standard oct summed voxel projection ( svp ) to use as an atlas to locate smaller volumes imaged using spfi , which is shown in fig1 ( a ). several volumes ( e . g ., 2 × 2 mm volumes ) can then be densely - sampled using spfi parameters ( e . g ., 2500 a - scans / frame ; 100 frames / volume ; 100 μs integration time , 25 s total imaging time ) at several locations across the macula ( see , e . g ., fig1 ( b ), 12 ( b ), and 12 ( e )), including landmarks such as the fovea ( see , e . g ., fig1 ( d )) and optic nerve ( see , e . g ., fig1 ( f ). given the sampling parameters , threshold frequency , and assuming a 20 μm scanning beam spot - size on the retina , the detectable velocity range for axially moving scatterers can be about 0 . 45 - 8 . 64 mm / s . parameters can be set for a lower velocity range , compared to flow phantoms , to adequately detect slow flow in small foveal vessels ( e . g ., & lt ; 30 mm / s ). the frequency window can be set such that the velocity - range has a particular fwhm ( e . g ., about 18 . 4 μm / s ) and is shifted at specified increments ( e . g ., about 82 . 8 μm / s ). bulk motion correction can be implemented prior to spfi windowing . velocity - resolved b - scans can be used to determine vascular size by fitting velocity profiles to parabolic flow curves and then calculating the zero - velocity crossing positions . finally , blood flow rates and total retinal flow can be calculated for a 2 × 2 mm velocity - resolved volume of human fovea . it is understood that spfi acquisition of the entire macula in a single volume dataset would be impractical . since velocity resolution and lateral sampling density are coupled in spfi by eq . ( 8 ), an order of magnitude increase in the lateral sampling density ( i . e ., a - scans / frame ) would be required . this requirement can limit the spfi sampling volume size by both the available memory in the acquisition software and total imaging time . depth - resolved vessel maps for each volume can be first reconstructed and overlaid onto structural oct data to distinguish macular vasculature ( see fig1 ( a ) through 12 ( e )). total blood flow measurements can then be calculated for a single 2 × 2 mm volume of the fovea ( see fig1 ( c )). the structural oct data can show the foveal pit and the associated spfi vascular map can confirm its location by resolving a circular avascular zone surrounded by a set of terminal capillaries . first , the resolved vessels in the volume can be identified as arteries and veins using the flow directionality information calculated by spfi ( see fig1 ( a )). vessel orientation and doppler angle can then be measured for all 17 resolved vessels , and velocity profiles can be measured at a single point for each vessel ( see , e . g ., fig1 ( a ), dots ). doppler angles can be measured by calculating the vessel cross - section displacement across sequential b - scans in the 3d datasets . the velocity profiles can then be fit to laminar flow curves to determine both the peak velocity and diameter of each vessel . finally , the vessel size , orientation , and velocity information can be used to calculated the total foveal blood flow . a representative b - scan ( see , e . g ., fig1 ( b )) taken across the foveal volume ( dotted line in fig1 ( a )) is shown with velocity - resolved flow content overlaid on top of the structural image . the velocity profiles and parabolic fits for each of the vessels are included ( see fig1 ( c ) through 13 ( i )) to show the strong correlation between velocities measured using spfi and their respective laminar flow velocity profiles ( mean r2 = 0 . 95 ). closer inspection of vessel 17 ( see fig1 ( i )) shows a blunted parabolic velocity profile , characteristic of red blood cell aggregation in microvasculature , consistent with rheological observations . a summary of the size , peak velocity , and flow measurements for all vessels identified in a representative 2 × 2 mm foveal volume are shown in table 1 below ( veins identified by (*)). in this representative volume , the smallest resolvable vessel was 13 . 64 μm , which is at the resolution limit of the retinal sdoct system . the average measured foveal vessel diameter is about 22 μm , and average arterial flow velocity is greater than average venous flow velocity , which is supported by similar measurements made using ldv . the detected velocities range from 5 . 97 - 30 . 22 mm / s , concurrent with human retinal vessels in the corresponding size range , also measured using ldv . finally , the total arterial and venous flow show a net inflow of blood into the fovea . this can be a result of the presence of unresolved veins in the volume and expected errors from measurements of vessel orientation angle and calculation of diameter , which would significantly impact blood flow calculations . while spfi allows for velocity - resolved volumetric blood flow imaging with velocity resolution comparable to that of doct , there are several differences between the modalities that can be noted . since velocity resolution and spatial oversampling are coupled , this allows for the flexibility to set sampling parameters to a desired velocity range . this flexibility would allow for improved data acquisition speeds over doct for imaging of small spatial volumes of moderately high flow velocities , but could potentially require longer scanning times for large scan areas with low flow . as discussed previously , the velocity resolving power of spfi falls off at the lower limits of velocity detection , as compared to doct , but is comparable for high flow velocities . since velocity - resolved spfi requires a sliding spatial frequency window , this increases the number of fourier transforms required by the number of velocity increments desired . however , this increase in computational complexity is essentially trivial since fft algorithms can be optimized and the transforms can be parallelized . finally , it has been shown that spatial frequency filtering results in a sensitivity improvement , which gives spfi a detection advantage over doct for small vessels at the resolution limit of the imaging system . the overall advantage of spfi over doct is in its improved sensitivity and customizability over velocity resolution , velocity range , and acquisition time in exchange for small increases in computational complexity . the present subject matter can be embodied in other forms without departure from the spirit and essential characteristics thereof . the embodiments described therefore are to be considered in all respects as illustrative and not restrictive . although the present subject matter has been described in terms of certain preferred embodiments , other embodiments that are apparent to those of ordinary skill in the art are also within the scope of the present subject matter .
the present subject matter relates to in vivo volumetric bidirectional blood flow imaging using single - pass flow imaging spectral domain optical coherence tomography . this technique uses a modified hilbert transform algorithm to separate moving and non - moving scatterers within a depth . the resulting reconstructed image maps the components of moving scatterers flowing into and out of the imaging axis onto opposite image halfplanes , enabling volumetric bidirectional flow mapping without manual segmentation .