full article imaging blood flow inside highly scattering ...€¦ · tivity and contrast unless a...

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FULL ARTICLE

Imaging blood flow inside highly scattering media usingultrasound modulated optical tomographyAltaf Hussain, Wiendelt Steenbergen, and Ivo M. Vellekoop*

Biomedical Photonic Imaging group, MIRA Institute for Biomedical Technology and Technical Medicine, University of Twente, PO Box217, 7500 AE Enschede, The Netherlands

Received 19 January 2017, revised 30 March 2017, accepted 18 April 2017

Keywords: medical and biological imaging, imaging through turbid media, acousto-optics, laser Doppler velocimetry

We report the use of ultrasound modulated optical to-mography (UOT) with heterodyne parallel detection tolocally sense and image blood flow deep inside a highlyscattering medium. We demonstrate that the UOT sig-nal is sensitive to the speed of the blood flow in the ul-trasound focus and present an analytical model that re-lates UOT signals to the optical properties (i. e.scattering coefficient, anisotropy, absorption, and flowspeed) of the blood and the background medium. Wefound an excellent agreement between the experimentaldata and the analytical model. By varying the in-tegration time of the camera in our setup, we were able

to spatially resolve blood flow in a scattering mediumwith a lateral resolution of 1.5 mm.

1. Introduction

The ability to measure blood flow in a noninvasivemanner with high sensitivity and accuracy is of fun-damental importance for studying a wide range ofbiological processes including local metabolism,neurovascular communication, pharmacodynamics,and stroke recovery. The commonly used imagingmodalities, such as MRI, CT and ultrasound (US)enable imaging of blood flow [1–3]. However, eachof these methods has its drawbacks: MRI is ex-pensive and has a low throughput; CT relies on ion-izing radiations; and ultrasound has a poor sensi-tivity and contrast unless a contrast agent is used [4].Optical methods, on the other hand, can potentiallybe cheap, safe, have high throughput and provide ul-timate sensitivity since a displacement in the orderof the wavelength of the light can potentially be de-

tected. Currently, optical techniques such as laserspeckle contrast analysis (LASCA), laser Dopplerand optical coherence tomography (OCT) are im-portant tools in the clinic [5–7]. However, thesetechniques are limited to the detection of superficialblood flow as the resolution deteriorates with depthdue to multiple scattering. Because of the high clin-ical relevance of blood flow measurements [8,9] in-tense research is being conducted to develop techni-ques with improved resolution, penetration depthand quantification ability [6,10–12]. Among thesetechniques, photoacoustics (PA) is being intenselyinvestigated for blood flow measurements due to itsadvantageous contrast mechanism based on absorp-tion [12]. In PA the blood flow is measured by meas-uring the Doppler shift in ultrasound wave (i. e. opti-cally generated). Similar to US flowmetry, however,the relatively long wavelength of the ultrasound

* Corresponding author: e-mail: [email protected]

J. Biophotonics 1–10 (2017) / DOI 10.1002/jbio.201700013

© 2017 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

J. Biophotonics 11: e201700013, 1 of 10 (2018) / DOI 10.1002/jbio.201700013

© 2017 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim © 2017 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

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makes the technique insensitive to small displace-ments.

Ultrasound modulated optical tomography(UOT) was originally developed to image absorbingobjects inside turbid media [13,14], and it has re-cently found exciting new applications such as focus-ing light into turbid media [15], measuring opticalfluence variations [16,17], and quantification of scat-tering inhomogeneities [18]. These applications ofUOT are mainly motivated by the invention of aheterodyne parallel speckle detection (HPSD)method using a CCD camera [19,20]. HPSD is supe-rior in terms of sensitivity compered to methodsbased on a single detector and allows for directquantitative measurements of ultrasonically modu-lated light [21]. Due to the superior sensitivity ofHPSD, the US burst that is used to modulate thelight can be as short as a few cycles that improvesthe axial resolution [22], further pulse inversion ap-proach is applied to improve the axial resolution[23]. However in our work we have only used shortUS bursts for axial resolution, which can be furtherimproved by combining it with pulse inversion.

Recently, the use of ultrasonically modulatedlight has been suggested for sensing the flow of scat-tering particles. Chandran et al. [24] demonstratedthat the temporal fluctuations in the reflected lightintensity can be linked to the flow speed of the mov-ing particles in the US focus. Tsalach et al. [25] pro-posed the use of ultrasonic modulation of light fordepth selective flow measurements. They showed us-ing simulation and experiment using single fast de-tector that the spectral width of the photocurrentfluctuation of ultrasonically modulated light is re-lated to the movement of scattering microspheres inthe ultrasound focus. However, the description byTsalach et al. of the dependence of UOT signal onthe flow of the optically scattering particles in theUS focus and dynamics of the background mediumis only qualitative. For truly quantitative estimationof flow using UOT, a robust detection method thatallows direct measurement of ultrasonically modu-lated light in absolute terms and an analytical modelthat describes the dependence of the UOT signal onoptical properties of the flowing particles and of thebackground medium is required.

Here we present the use of UOT with HPSD forimaging depth-resolved blood flow inside highlyscattering media. We experimentally show the de-pendence of the HPSD signal on the speed of bloodflowing through a tube embedded in a scatteringmedium. Further we present an analytical modelthat relates the HPSD signal to the optical proper-ties (i. e. scattering coefficient, anisotropy, absorp-tion, and flow speed) of moving particles and thebackground medium. Finally, we demonstrate a sim-

ple flow imaging method that relies on comparingUOT images taken with a short and a long in-tegration time of the CCD. In comparison with workof Chandran et al. and Tsalach et al. [24, 25] on sens-ing flow of scattering particle using UOTusing a sin-gle, fast photodetectors, we employ a parallel de-tection scheme with a CCD camera. Our methodrelates to these studies as LASCA relates to laserDoppler flowmetry [26].

2. Material and methods

2.1 Phantom

We performed experiments on a slab of scattering tis-sue phantom of dimensions 33331.5 cm3 that con-tained a polyethylene tube mimicking a blood vessel.The acoustic impedance of polyethylene is 1.73 MRa-yls [27] which is close to the acoustic impedance of wa-ter 1.5 MRayls, leading to 92% acoustic transmissionwhen placed in water. The tube had an inner and out-er diameter of 0.96 and 1.28 mm respectively. For ex-periment 1 and 2, the configuration of the sample usedis illustrated in Figure 1b, where the tube is placed5 mm under the surface that is illuminated. For experi-ment 3 the configuration of the sample used is illus-trated in Figure 1c, where the depth of the tube underthe surface at the first and second pass is 5 and 10 mmrespectively. The tissue phantom with a reduced scat-tering coefficient (m

0s) of approximately 0.7 mm�1 was

prepared by mixing 7 % (w/v) of 20 % Intralipid (Sig-ma Aldrich) and 3 % (w/v) of Agar (Sigma Aldrich)in 90 ml of water. Blood was obtained from an anony-mous healthy volunteer organized by the TNW-ECTM-donor service at the University of Twente. Thetotal hemoglobin concentration of the blood sampleswas 15 g/dl and the oxygen saturation was 96 %. Oxy-gen saturation and total hemoglobin of the blood sam-ples were measured using a blood oximeter (Avoxim-eter 1000E, ITC point of care) before and after theexperiment for comparison and were found to remainunchanged over the duration of the measurement.Blood was pumped through the nylon tube in a closedsystem using two syringe pumps (Figure 1b and 1c),one pushing and the other withdrawing at the samerate, this configuration also prevented the blood frombeing exposed to air during the experiment.

2.2 Heterodyne parallel speckle detectionmethod for ultrasound modulatedoptical tomography

In UOT, the medium is illuminated with coherentlight of frequency wo, and a region of interest

2 A. Hussain et al.: Imaging blood flow inside highly scattering media

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(ROI) is insonified with focused ultrasound (US)of frequency wUS (see Figure 1a). The ROI can beplaced anywhere inside the medium thanks to thelow scattering of ultrasound. A fraction of the dif-fused light interacts with the ultrasound and getsmodulated at �wUS. In effect, the insonified re-gion becomes a source of ultrasonically modulatedlight of frequency wo�wUS. The detection of theultrasonically modulated light at the mediumboundary was done using heterodyne parallelspeckle detection (HPSD) [28]. Our setup, shownin Figure 1a, consists of a CCD camera, an UStransducer and a Mach-Zehnder interferometer.The light in the reference arm reaches the CCDcamera at an angle, so that we can perform off-axis holography. The reference arm angle is ad-justed by rotating the beam splitter (BS) in frontof the CCD camera (see Figure 1). The BS is fixedon a manual rotation stage and the angle of refer-ence arm is chosen such that the hologram of theaperture (i. e. heterodyne term) is off axis enoughso that it is isolated from the on axis background.The alignment is done by looking at the real timehologram of the aperture. Additionally, acousto-optic modulators (AOM) are used to shift the fre-quency of the light in the reference arm to wo+wUS, in order to match the frequency of the modu-lated light emanating from the sample.

The CCD camera records the image of the in-

tensity Ið~rÞ ¼ RT0

Eð~r; tÞE*ð~r; tÞdt, where T is the in-

tegration time, E(r,t) is the total electric field, andE*(r,t) is its complex conjugate. The electric field is asum of the field of the reference beam, the diffuse

light modulated at frequency wo+wUS, the non-modulated diffuse light, and light modulated at wo-wUS or higher ordered frequencies. Of these con-tributions, only the first two interfere with each oth-er while the others only contribute to a backgroundsignal Ibgð~rÞ. Therefore, we can express the recordedintensity at the CCD as

Ið~rÞ ¼ Ibgð~rÞþZT

0

E*USð~r; tÞELOei~kLO�~r þ EUSð~r; tÞE*LOe�i~kLO�~r

h idt

ð1Þ

where ELO is the electric field amplitude of the lightfrom the reference arm, ~kLO is the wave vector ofthe reference beam,~r is the position vector and EUS

is the electric field amplitude of the ultrasonicallymodulated light. The time dependence of the fieldamplitude EUS is due to the temporal variations inthe field that are caused by moving particles.

The camera image (Figure 2a) contains fringesthat correspond to the heterodyne term E*LOEUSð~r; tÞ.One can isolate the heterodyne term from the back-ground light intensity Ibgð~rÞ in k-space after spatiallyFourier transforming the CCD image [21], a methodknown as off-axis holography. In k-space Ibg is cen-tered at the origin, and its spread is defined by thespeckle size relative to the size of the camera pixels.The position of the heterodyne term is defined bythe frequency of the fringes (Figure 2a). In turn, thisfringe frequency depends on the angle of the wavevector (~kLO) of the reference beam relative to thenormal of the image sensor. The angle of the refer-

Figure 1 (a) schematic of the ultrasound modulated tomography setup based on heterodyne parallel speckle detection(HPSD). AOM: acousto-optic modulator, BS: beam splitter, CCD: charge coupled device camera. Two syringe pumps areused for enforcing a controlled blood flow, (b) schematic of the sample used in experiment 1 and 2, (c) schematic of thesample used in experiment 3.

J. Biophotonics (2017) 3

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ence beam is chosen such that the heterodyne termis separated from the Ibg term, and still falls withinthe k-space image completely (Figure 2b). As a re-sult, HPSD enables measurement of E*LOEUSð~r; tÞ inisolation from the diffuse background light emanat-ing from sample, written as

HðrÞ ¼ E*LO

ZT

0

EUSð~r; tÞdt ð2Þ

where is T is the integration time of the CCD cam-era. In case the medium is dynamic, the integrationtime of the camera suppress the interference fringes(see Figure 2c), hence reducing heterodyne term.

Further, we define the modulation depth as,

M � Hj j2� �ELOj j2� �

ELj j2� �T2

ð3Þ

where hi denotes averaging over CCD pixels, T jEL j2 is the CCD signal corresponding to the totalamount of diffuse light emanating from the sampleand it is measured by switching off the referencebeam. T jELO j 2 is the CCD signal corresponding tothe light in the reference beam and it is measuredthe while the sample arm is switched off. Using Eq.

2, the expression for the modulation depth M (i. e.Eq. 3) can be written as,

M ¼ 1ELj j2� � h

ZT

0

ZT

0

EUSð~r; tÞE*USð~r; tÞdtdti

¼ 2 EUSj j2� �ELj j2� �

T

ZT

0

ð1� t=TÞCEðtÞdtð4Þ

where, CEðtÞ � EUSð~r; tÞE*USð~r; tÞ� ��

EUSð~r; tÞj j2� �is

the field autocorrelation function and t is the timewith respect to the start of acquisition of the de-tector. We arrived at Eq. 4 by assuming that EUS isergodic and CE is an even function of t (see Ref. [29]for a similar derivation involving the intensity corre-lation function). Eq. 4 shows that the modulationdepth M in HPSD is directly linked with the fieldautocorrelation function CE of the ultrasonicallymodulated light. In areas with more flow, CE will de-crease as a function of time more rapidly. Therefore,we expect the modulation depth to decrease whenthe ultrasound focus is placed in areas with a higherblood flow.

2.3 Experimental setup

Details of the setup and the experiment are as fol-lows. The sample is illuminated by a 760 nm CW la-ser (Coherent MBR EL Ti:sapphire, pumped by aCoherent Verdi 6), beam diameter 3 mm and ultra-sonically modulated light is detected with a CCDcamera (GRAS-14S5M�C, 138431036) through anaperture of diameter 3 mm. The CW laser beam ischopped into pulses of 1 ms at repetition rate of25 kHZ using AOM1. Laser power incident on thesample was measured to be 18 mW, which leads tothe optical irradiance of 0.25 W/cm2. The laser puls-es are synchronized with ultrasound bursts of 1 ms toachieve a spatial resolution of 1.5 mm along the USpropagation axis. The input laser beam is split intoreference and sample arms using a beam splitter(BS). The reference arm is shifted in frequency by5 MHz using AOM up (+80 MHz) and AOM down(�75 MHz) to match the frequency of the US trans-ducer. The US transducer (Panametrics V309, diam-eter of 12.7 mm and focal length of 20.3 mm) wasdriven by a 5 MHz signal amplified using a poweramplifier (ENI 350 L). Short bursts (0.6 ms) of fiveUS cycles were used for local modulation of light re-sulting in a pressure of approximately 1 MPa in theUS focal zone. The delays between the US burst andthe laser pulse were set using a function generator,which allowed the ultrasound to propagate (z-axis) a

Figure 2 illustration of the principle of the HPSD methodfor sensing movement of scattering particles, the whitebox marks the position of heterodyne term in k-space. a)CCD camera image with integration time less than the de-correlation time of the medium, b) Fourier transform of a,c) CCD camera image with integration time longer thandecorrelation time of medium, d) Fourier transform of c.





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configurable distance before interrogating the re-gion of interest with light.

2.4 Analytical model for blood flowmeasurements using ultrasoundmodulated optical tomography

To describe our blood flow measurements quantita-tively, we apply the following physical model. Weconsider an external light source S that is 100 % cor-related. The correlated light propagates from thesource at~ro to the US focus at~r with the propagatorGð~r;~ro; tÞ, where the light gets tagged with a taggingefficiency of h(r). The ultrasonically tagged lightpropagates further to a point ~rd at the boundary ofthe medium with the propagator Gð~rd;~r; tÞ where itgets detected. The amount of ultrasonically taggedcorrelated light at point~rdis given as,

Chð~rd; tÞ ¼ZZ

Gð~rd;~r; tÞhð~rÞGð~r;~ro; tÞSð~roÞdrodr ð5Þ

We describe the propagation of correlated lightwith the following three dimensional inhomoge-neous diffusion equation [30],

Dr2Gð~r;~ro; tÞ � ðmdðtÞ þ maÞGð~r;~ro; tÞ ¼ �dð~r �~roÞ ð6Þ

where md is the decorrelation coefficient, which is afunction of time difference t. ~r is the three-dimen-sional position vector, D is the diffusion coefficientin meter and ma is the absorption coefficient. Duringthe propagation, light may encounter moving par-ticles that decorrelate the light to some extent. Sincethe decorrelated light is not detected in our hetero-dyne setup (as described above), we consider it lost.Therefore, we can model the decorrelation as havingthe same effect on light as absorption. The decorre-lation coefficient md is defined as the number of scat-tering events per propagated meter (ms) times thefraction of light that gets decorrelated at each scat-tering event. In line with [30] we define,md ¼ msð1� e�2<Dr2>k2oð1�gÞÞ, where g is the anisotropyfactor and <Dr2> is the mean squared displace-ment of the moving particles, and ko�2pn/l where nis the refractive index and l is the wavelength of thelight. In our experiments, we have blood flowingthrough a nylon tube, we treat it as a directed flowsuch that Dr ¼ vt, with v the flow speed of theblood, and t the time with respect to the start of theacquisition time of the detector.

To model the experiment in an analytically tract-able way, we model the scattering medium as a cyl-inder of radius ro, with a tube of radius rT with bloodflowing through it placed at r=0 (see Figure 3). As-

suming radial symmetry, the diffusion equation incylindrical coordinates reads

@2Gðr;ro ;tÞ@r2 þ @Gðr;ro ;tÞ

r@r � mdðtÞþmaD Gðr; ro; tÞ

¼ �dðr�roÞD

ð7Þ

where r is the radial coordinate. Further we define,

a ¼ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiðmd;inþma;inÞ

Din

qand b ¼

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffimd;outþma;out

Dout

q, where

Din ¼ 3ðm0s;in þ md;in þ ma;inÞ

� ��1 is the diffusion co-efficient inside the tube with the blood flow andDout ¼ 3ðm0

s;out þ md;out þ ma;outÞ� ��1 is the diffusion co-

efficient of the background medium in which thetube is placed. The solution to Eq. 7 is,

Gðr; ro; tÞ ¼bKoðbrÞ � roI0ðbrÞK0ðbroÞ

DoutrT< r < ro

aI0ðarÞ 0 <¼ r < rT

8><>: ð8Þ

Here, Io and Ko are the Bessel functions of the firstand second kind respectively, coefficients a and b aretime dependent because of their dependence on md

and will be solved later. In Eq. 8 we chose ro>rT andomitted the terms that diverge as r!1 or r!0.

Next we solve for coefficients a and b; consider-ing the continuity of intensity Gðr; ro; tÞand the flux(�D @Gðr;ro ;tÞ

@r ) at the tube surface (r= rT),

a ¼ � r0K0 r0bð ÞaDinrTI1 rTað ÞK0 rTbð Þ þ bDoutrTI0 rTað ÞK1 rTbð Þ

b ¼ r0K0 r0bð Þ aDinI1 rTað ÞI0 rTbð Þ � bDoutI0 rTað ÞI1 rTbð Þð ÞDout aDinI1 rTað ÞK0 rTbð Þ þ bDoutI0 rTað ÞK1 rTbð Þð Þ

ð9Þ

Here, a, and Din are both time dependentthrough their dependence on md. We can now useEq. 5 to calculate the fraction of correlated light that

Figure 3 illustration of the model geometry. rT is the tuberadius, ro is the radius of the object and the location of thesource, rUS and DZ are the radius and length of the USfocus.





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propagates from the source, interacts with the USfocus and reaches the detector. We modeled the ul-trasound focus as a cylinder segment of radius rusand height DZ concentric with the tube in whichblood flows (see Figure 3), where part of the US fo-cus lies outside the tube containing blood flow. Atthe ultrasound focus, the correlated light gets tag-ged, and locally acts as a source of frequency shiftedcorrelated light. This tagged light propagates out-wards from the ultrasound focus, and results in a netoutgoing flux of frequency shifted correlated light.As a further simplification, we assume that the de-tector is placed at the same distance from the tubeas the source (rd= ro), such that there is a symmetryin the propagation of inward flux from the sourcetowards the ultrasound focus, and outward flux to-wards the detector. Using Eq. 5, the ultrasonicallytagged correlated light that arrives at the detector isgiven as,

Chðr; tÞ ¼ 2pDZhZrUS

0

G2ðr; ro; tÞrdr ð10Þ

Considering the definition of the modulationdepth given in Eq. 4, the expression for the modu-lation depth based on the model is given as,

M ¼ 1T

ZT

0

ð1� t

TÞZrUS

0

G2ðr; ro; tÞrdrdt ð11Þ

We evaluated the Eq. 11 numerically to estimatethe modulation depth by using the following param-eters ro=6 mm, rUS=0.5 mm, DZ=1.5 mm, rT=0.48 mm, l=760 nm, m

0s;out=0.7 mm�1,

m0s;in ¼ a l

500ðnmÞ

� bwhere, l is the wavelength in nm,

a=220 mm�1 and b=0.66 for the whole blood [31],n=1.33, and g=0.99 matching with the ex-perimental situation and we fitted the proportion-ality constant to account for unknown tagging effi-ciency h and the total untagged light at the detector.

The model assumes a block flow of scatteringparticles in the tube whereas in reality the flow pro-file of blood is known to be parabolic. This flow pro-file reduces the effective radius of the tube in whichthe blood flows. We found that an effective tube ra-dius of rT-eff=0.7rT matches well with the ex-perimental data.

3. Results

To demonstrate the concept of depth-resolved meas-urements of blood flow, we performed three experi-

ments. In the first experiment, we investigated therelation between the ultrasound modulation depthand the acquisition time of the CCD camera for afixed flow speed. We kept the flow speed of theblood fixed at 0.2 mm/s through the tube and variedthe integration time of the CCD camera. Multipleline scans were made along the propagation direc-tion of the US for integration times varying from1.75 ms to 4 ms with a step of 0.25 ms. Here, 1.75 mscorresponds to the shortest integration time that stillgave enough signal to be interpreted meaningfully.

Figure 4a shows the line scans along the z-axis,each line scan is acquired with a different in-tegration time of the CCD camera. In Figure 4a, wecompare two situations, one where the US focusoverlaps the region with blood flow (at x=5 mm),and one where the US focus is placed in the staticbackground (at x=4 mm). The modulation depthprofile in the static background case is maximumwhen the US burst is centered at z=21.75 mm. Thisis because that the probability of light to propagatefrom illumination aperture to the US burst and getmodulated and propagate further to the detectionaperture is maximum when the US burst is in theplane of illumination and detection aperture whichis at z=21.75 mm. The line scans at x=5 mm in Fig-ure 4a, correspond to the case where the US burstpropagate through the tube with constant blood flowplaced. Therefore, in these line scans we observe alocal minima at the location of the tube (i.e z=22 mm). This decrease in modulation depth at thelocation of tube is the result of the decorrelation ofultrasonically modulated light by the blood flow. Itis apparent from Figure 4a that the modulationdepth at these minima decreases with an increase inintegration time. Figure 4b, shows the measuredmodulation depth and modulation depth predictedby the model as a function of the integration time ofthe CCD camera when the US focus was over-lapping with the tube. The error bars on the ex-perimental data are stand deviation of five measure-ments. The five measurement are performed aftereach other for a fixed integration time of the CCD,then the integration time of the CCD is changed andthe next measurements are performed. Both themodel and the experimental data show that themodulation depth decreases with increasing in-tegration time of the CCD camera. The comparisonbetween the model and the experiment data for aneffective radius of the tube with blood is considered0.7 times the geometric radius of the tube yielded R-Squared value of 0.92.

In the second experiment, we investigated the re-lation between the modulation depth and the localflow speed of the blood in the turbid medium. Wekept the integration time of the CCD camera fixed





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at 2.0 ms and changed the flow speed of the bloodfrom 0 mm/s to 2.75 mm/s with increments of0.23 mm/s per measurement. The delay time be-tween US burst and light pulse for these measure-ments was such that the US focus overlapped withthe tube. Figure 5 shows the measured modulationdepth and modulation depth predicted by the modelas a function of the flow speed of the blood. Withouttaking into account the data point corresponding toflow speed 0 mm/s the comparison between the ex-

perimental data and the model yielded R-squaredvalue of 0.87. In case of flow speed 0 mm/s the mod-ulation depth predicted by model is factor of twohigher than the experimental value. This discrepancyat a flow speed of 0 mm/s may be due to the factthat optical properties of the blood change as bloodstops flowing [32], this aspect will be discussed fur-ther in the discussion session.

In the third experiment, we explored the feasibilityof the technique to make depth-resolved maps ofblood perfusion in turbid media. We scanned a regionof 4311 mm2 in the scattering medium containing anylon tube passed twice through the medium (see Fig-ure 1c). The depth of the tube under the surface at thefirst and second pass was 5 and 10 mm respectively.The blood was pumped through the tube at flow speedof 0.2 mm/s. The medium was scanned with the US fo-cus in the xz-plane, the scan along the x-axis was ach-ieved mechanically while the z-axis was scanned bydelaying laser pulses relative to US bursts. Two scanswere performed consecutively, for the first scan the in-tegration time of the camera was 1.5 ms (scan A) andfor the second scan the integration time was 10 ms(scan B). Then the ratio (MA/MB) of the modulationdepths in the two scans is calculated. Under ideal cir-cumstances, in regions with blood flow the ratio willexceed one. Everywhere else the ratio will be unity,since the variation in modulation depth due to absorp-tion is independent of the integration time of the CCDcamera.

Figure 6a and b, show the two scans (A and B) inthe xz-plane with an integration time of 1.5 ms and

Figure 4 (a) UOT line scans through the tube (x=5 mm) and before the tube (x=4 mm) along the z-axis with varying in-tegration time, (b) ultrasound modulation depth versus integration time of the CCD camera, for a flow speed of 0.2 mm/s ofblood through the tube. The tube is placed in the US focus (z=22 mm, with z=0 being the surface of the transducer) and5 mm (along x-axis) below the surface of the phantom. rT is the actual radius of the tube, whereas rT-eff is the effective radiusused in the model.

Figure 5Measured modulation depth as a function of flowspeed of blood in the tube, for a fixed integration time of2 ms. The tube is placed in the US focus (Z=22 mm) and5 mm (along x-axis) below the surface of the phantom.





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10 ms respectively. In both scans the regions close tothe illumination and detection apertures show a highmodulation depth which decreases towards the cen-ter of the sample. This is because, when the US fo-cus is close to illumination/detection aperture it ismore probable that the detected light has interactedwith the US focus. In scan B the drop in modulationdepth at the location of tubes is larger than in scanA, and it is caused by a combination of decorrela-tion and absorption of light by blood in the tubes.From scan A or scan B alone, it cannot be de-termined whether the drop in signal is caused by ab-sorption or by blood flow, not is it possible to re-solve the two tubes spatially.

Figure 6c shows the ratio of modulation depthsof scans A and B. As expected, the ratio peaks atthe centers of the two tubes, clearly demonstratingthe ability to form depth-resolved images of theblood flow. Figure 6d shows the comparison of theline profiles of the conventional UOT scans A, Band the ratio image along the x-axis for z=22 mm.The two peaks in the line profile (blue solid line, d)of the ratio image correspond to the two tubes withblood flow. Based on the full width half maximum ofthese two peaks, indicated by vertical doted lines inFigure 6d, we estimated the lateral along the z-axisis defined by the length of the US burst which is

equal to 1.5 mm in our experiment, given the USburst length of 1 mS and speed of sound 1500 m/S inwater. The ratio image provides a depth-resolvedsemi-quantitative map of blood flow inside the tur-bid medium. Semi-quantitative here means that thesignals are directly related to the flow speed ofblood, given that the medium is acoustically homo-geneous such that the shape and size of US focus re-mains unaltered while scanning a region of interest.The flow profile in c is slightly elongated along theoptical axis (x-axis). This distortion may be causedby the highly forward scattering property of thephantom and blood in combination with a trans-mission mode experimental geometry.

4. Discussion and conclusion

We presented the use of ultrasound modulated opticaltomography based on heterodyne parallel speckle de-tection to make semi-quantitative depth-resolved im-ages of blood flow inside the scattering media. Weshowed theoretically that the modulation depth is di-rectly linked to the field autocorrelation function ofthe ultrasonically modulated light field EUS and mod-eled the propagation of the field autocorrelation func-tion of ultrasonically modulated light to better under-

Figure 6UOT scans of turbid media having two nylon tubes embedded at x=3 and 8 mm relative to scan area with bloodflow. a) scan A: modulation depth, integration time 1.5 ms, b) scan B: modulation depth, integration time 10 ms c) perfu-sion image: ratio of modulation depths (scan A and B,) d) comparison of line profiles along x-axis at z=22 mm for twoscans A (dotted black) and B (dashed red) and ratio image (solid blue). The tubes with blood are located at x=3 mm andx=8 mm.





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stand the relation between the physical properties ofmoving particles and the measured modulation depth.Experimentally, we measured the modulation depth asa function of acquisition time of the CCD camera fora fixed flow speed of blood in the US focus and com-pared the results with the prediction of the model. Weobserved a good agreement between the model andthe experimental data using effective radius of0.7 times the actual tube radius to account for flowprofile.

The observed decrease in modulation depth withincreasing acquisition time is analogous to the LAS-CA signals, where the speckle contrast decreaseswith the integration time of the CCD camera [33].Our findings suggest that our technique is similar toLASCA with the added benefits of depth selectivity.Further we observed that both the model and theexperimental show that the modulation depth de-creases with increasing speed of blood flow for a giv-en integration time of the CCD camera. However,at a flow speed of 0 mm/s, the modulation depth pre-dicted by the model is two times higher than ex-perimentally measured modulation depth. One pos-sible reason for this discrepancy might be the factthat the model assumes the same optical propertiesof blood at all flow speeds, while it has been demon-strated that the scattering of blood decreases as flowspeed approaches zero due to Rouleaux formation[32]. The decrease in scattering in turn reduces thetagging efficiency of the US thereby decreasing themeasured modulation depth, since it is dependenton the scattering coefficient of the medium [34]. Al-though our observation is consistent with the phe-nomenon, the hypothesis requires further inves-tigation and quantification, either by quantifying thechanges in optical properties at lower flow speed ofblood and taking that into account in the model orby simulating the scattering fluid that resemblesblood and its optical properties are not dependenton the flow speed. Our future work will focus on thisaspect and also on quantifying the performance ofUOT as a technique to measure blood flow in com-parison with ultrasound Doppler and photoacousticflowmetery in terms minimum/maximum measure-able flow speeds and sensitivity.

The minimum flow speed that can be detected bythis technique is only limited by the dynamics of thebackground medium. By increasing the integrationtime of the CCD, the system may be made sensitiveto even the slowest movements inside the sample,with the caveat that red blood cells aggregate atsuch low flow speeds. Considering our results of ex-periment 2 we were able to accurately measure theminimum flow speed of 0.3 mm/s, which is 20 timesmore sensitive than the ultrasound power Dopplerusing a transducer with a center frequency of 6 MHz

[35]. The minimum detectable flow speed usingpower Doppler is dependent on the frequency of ul-trasound transducer used which puts physical limiton minimum detectable flow speed and imagingdepth [35]. In photoacoustic flowmetery the contrastis superior to ultrasound however the detection prin-ciple is similar in nature to that of ultrasound, there-fore the minimum detectable flow speed limit is ex-pected to be identical.

The maximum measurable flow speed is limitedby the minimum integration time required to meas-ure a signal on the CCD. By employing a differentdetection technique with tandem pulses designed foracousto-optic imaging [36], the effective integrationtime may be reduced to as low as 20 ns, which is 4orders of magnitude shorter than the integrationtime used in our experiment number 2. In experi-ment number 2 we have seen that at the integrationtime of 2 ms the flow speed of 2.75 mm/s results inmaximum measureable decrease in modulationdepth, as a result the system becomes insensitive toflow speed higher than 2.75 mm/s. However, whenusing integration time of 20 ns the limit of maximummeasureable decrease in modulation depth wouldoccur at 27.5 m/s, making maximum measurableflow speed 4 orders of magnitude higher than in ourcurrent set up.

Our findings suggest that UOTwith HPSD has apotential for depth-resolved blood flow imaging inhighly diffusive media. Our model shows that thedecorrelation rate of the background medium sur-rounding the blood vessel affects the quantitative re-sults. Therefore, we expect that the ability of UOTfor quantitative imaging can be further improved bycombining it with LASCA to obtain the global de-correlation coefficient md of the background me-dium, and using this measured coefficient in ourmodel to compensate for the background dynamics.In diffusive media where the light trajectories arescrambled enough to form a near isotropic light irra-diance, the situation similar to presented in experi-ment 3, the spatial resolution of the technique is ap-proximated to be 1.5 mm. We expect the techniquesability to quantitatively measure flow speed of bloodin two blood vessels would be compromised if thetwo blood vessels are spatially unresolvable. Oneother perceivable limitation may occur when a smallblood vessel is situated close to large blood vesselwith a high-speed blood flow in it. In this case, thelarger blood vessel will cause a shadow since most ofthe light in its surrounding would have been decor-related through its interaction with it and will affectthe visibility of the small blood vessel. The quantifi-cation of such effects is outside the scope of thisproof of principle paper and will be investigated inour future work.





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In conclusion, we presented a method for depthresolved imaging of blood in highly scattering mediabased on UOT using HPSD method. We demon-strated through analytical model and experimentthat UOT signal is dependent on the flow speed ofblood in the US focus for a given integration time ofthe detector and vice versa. We demonstratedthrough a simple experiment that the ratio of thetwo UOT scans performed using different in-tegration time of the CCD camera can be used to dodepth resolved imaging of blood vessels inside thescattering media.

We acknowledge Steffen Resink, Berkan Baseland Robert Molenaar for their input and valuablediscussions. This research is supported by the Tech-nology Foundation in the Netherlands STW, underDemonstrator grant, project 14546 and the Euro-pean Research Council under the European Union’sHorizon 2020 Program / ERC Grant Agreement no678919, and by the MIRA Institute of the Universityof Twente.

Author biographies Please see Supporting Informationonline.

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