A thin-walled carotid vessel phantom for Doppler ultrasound flow studies

Ultrasound in Med. & Biol., Vol. 30, No. 8, pp. 1067-1078, 2004Copyright © 2004 World Federation for Ultrasound in Medicine & Biology

Printed in the USA. All rights reserved0301-5629/04/$–see front matter

doi:10.1016/j.ultrasmedbio.2004.06.003

● Original Contribution

A THIN-WALLED CAROTID VESSEL PHANTOM FOR DOPPLERULTRASOUND FLOW STUDIES

TAMIE L. POEPPING,*† HRISTO N. NIKOLOV,* M EGHAN L. THORNE*† andDAVID W. HOLDSWORTH*†‡

*Robarts Research Institute, London, ON, Canada; and Departments of†Medical Biophysics; and‡DiagnosticRadiology and Nuclear Medicine, University of Western Ontario, London, ON, Canada

(Received 8 August 2003;revised 29 May 2004; in final form 3 June 2004)

Abstract—A technique is discussed for producing a robust ultrasound (US)-compatible flow phantom thatconsists of a thin-walled silicone-elastomer vessel with a lumen of arbitrary geometry, embedded in an agar-basedtissue-mimicking material (TMM). The TMM has an acoustic attenuation of 0.56 dB cm�1 MHz�1 at 5 MHz,with nearly linear frequency-dependence and acoustic velocity of 1539� 4 m s�1. The vessel-mimicking material(VMM) has an acoustic attenuation of 3.5 dB cm�1 MHz�1 with linear frequency-dependence and an acousticvelocity of 1020� 20 m s�1. Scattering particles, which are added to the VMM to increase echogenicity and addspeckle texture, lead to higher attenuation, depending on particle concentration and frequency. The VMM isstable over time, with a Young’s elastic modulus of 1.3 to 1.7 MPa for strains of up to 10%, which mimics humanarteries under typical physiological conditions. The phantom is sealed to prevent TMM exposure to air or water,to avoid changes to the acoustic velocity. © 2004 World Federation for Ultrasound in Medicine & Biology.

Key Words: Flow phantom, Vessel-mimicking material, Tissue-mimicking material, Carotid artery bifurcation,Doppler ultrasound, Sylgard 184, Silicone elastomer, Elastic modulus, Acoustic velocity, Speed of sound, Acousticattenuation.

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INTRODUCTION

Flow phantoms are an essential tool for the developof Doppler ultrasound (US) techniques for investigablood flow. To produce realistic flow conditions, ithighly desirable to be able to produce arbitrary geotries, including branching and stenosis, that closelysemble or mimicin vivo geometries. Typically, wall-lesvessels have been used to avoid mismatch problemacoustic properties (Guo and Fenster 1996; PattersoFoster 1983; Rickey et al. 1995), and to produce mcomplicated geometries by using a lost-material catechnique (Poepping et al. 2002; Smith et al. 19However, with wall-less vessels, the tissue-mimickmaterial (TMM) may be vulnerable to changes dueexposure to air or water (Hoskins and Ramnarine 20incompatibilities with the blood-mimicking fluid (BMF

Dr. Poepping’s current address is Medical Physics, UniversEdinburgh, Chancellor’s Building, 49 Little France Crescent, Eburgh, EH16 4SB, UK.

Address correspondence to: D.W. Holdsworth, Imagingsearch Laboratories, Robarts Research Institute, P.O. Box 501Perth Drive, London, Ontario N6A 5K8 Canada.E-mail: david
[email protected]
1067

or breakage and leakage problems, as addressed bynarine et al. (2001). Hence, it is also desirable to hacompletely sealed phantom, where the agar-basedis sealed from air or fluid exposure.

A number of flow phantoms are described inliterature and have been reviewed by Hoskins (19However, most of the flow phantoms previously deoped have consisted of a simple geometry, sucstraight tubes, or a stenosed geometry formed by shing tubes (McDicken 1986), inserting tapers (CalilRoberts 1985; Hoskins 1997) or extracting rods (and Fenster 1996; Patterson and Foster 1983; Rickal. 1995). The development of more realistic phangeometries has been primarily limited by the lackversatile vessel-mimicking materials (VMMs).

A primary consideration for the development onew VMM was the ability to fabricate a robust flophantom, with vessels of anatomically correct modwhich can withstand physiological flow stresses.main design considerations include robustness, stabelasticity, wall thickness, speed of sound, acoustic auation, echogenicity and, finally, practical considerat
for fabrication, such as versatility, variations in echoge-

1068 Ultrasound in Medicine and Biology Volume 30, Number 8, 2004

nicity and the incorporation of plaques or irregular sur-faces.

In addition to simulating anatomical geometries, themain physiological features to be included are the wallthickness and elasticity. It is desirable to match the wallthickness for realistic imaging and to match the elasticitybecause of its contribution to the damping and propaga-tion of the pulsatile waveform during flow. Studies haveshown that the wall thickness of the carotid arteries isapproximately 0.65 to 0.95 mm, depending on gender,race and duration of diabetes (Bonithon-Kopp et al.1996; Folsom et al. 1994; Riley et al. 1992; Wagen-knecht et al. 1997). Also, arterial walls have been shownto have a nonlinear elastic behaviour (Roach and Burton1957), due to the different elasticities of the elastin andcollagen components. The Young’s elastic modulus ofelastin, 0.3 MPa (Caro et al. 1978), is much lower thanthat for collagen, 300 to 2500 MPa (Dobrin 1978), andthe resulting arterial elasticity is similar to elastin fornormal physiological strains, and approaches that of col-lagen for much higher strains. The incremental Young’selastic modulus in the circumferential direction has beenfound, in normal human abdominal aorta sections, torange from approximately 0.04 to 2 MPa for correspond-ing transmural pressures of 0 to 90 mmHg (0 to 12 kPa)(Drangova et al. 1993) and, in the human iliac artery, torange from approximately 0.05 to 1.45 MPa for trans-mural pressures of 0 to 120 mmHg (0 to 16 kPa) (Ryanand Foster 1997).

The important acoustic properties to be matched arethe speed of sound (SoS) and the acoustic attenuation.Ideally, a VMM should be acoustically well matched toboth the BMF and the TMM. A mismatch in the speed ofsound between the VMM and surrounding material(BMF or TMM) will produce reflection and refraction ofthe US beam, which can result in distortion of the Dopp-ler spectrum (Hoskins 1994; Hoskins and Ramnarine2000; Thompson et al. 1990). Few suitable materialshave been found that match both the acoustic attenuationand the SoS of soft tissue (0.5 to 0.7 dB cm�1 MHz�1

and 1540 m s�1, respectively). Many materials havebeen investigated and described for use as the VMM inflow phantoms, as seen in the review by Law et al.(1989). C-flex� (Cole-Parmer, Vernon Hills, NY) is onepopular choice for a VMM, with a speed of sound (1553m s�1) similar to that of tissue, but with a much higherattenuation (28 dB cm�1 at 5 MHz) (Hoskins 1994) thatvaries with frequency to the power of 1.8 (Ramnarine etal. 2001). Unfortunately, most of the materials suggestedas a VMM are generally available only as ready-madetubing, such as C-flex� and latex, and are, thus, limitedto straight-tube or stenosed-tube geometries with a man-ufacturer-specified wall thickness and fixed echogenic-
ity.
Silicone allows the opportunity to produce vesselswith arbitrary wall thickness and lumen geometry, due toits availability in a liquid form that can be cast to form aflexible transparent elastomer. These elastomers arestrong, yet compliant, chemically stable over time, anddo not absorb a significant amount of water (unlike latexrubber). Silicone has been used previously to create tubesof various thicknesses for wave propagation experiments(Pythoud et al. 1994), stenosed tubes for laser visualiza-tion (Cao and Rittgers 1998) and branched vessel modelsfor flow visualization (Melbin et al. 1982), by paintinglayers of an elastomer (Sylgard� 184) onto a rod or cast.A casting technique has also been used with this materialto produce a branched carotid model for use with laserDoppler velocimetry and flow visualization (Secomski etal. 2003).

In what follows, the fabrication of a sealed, US-compatible, flow test object that consists of a thin-walledsilicone-elastomer vessel of arbitrary geometry embed-ded in a TMM is discussed.

METHODS

MaterialsThe VMM was comprised of a silicone (polydim-

ethylsiloxane) elastomer (Sylgard� 184, Dow CorningCorp., Midland, MI, USA; density 1.05 g/cm3) formedby combining a base and a curing agent in a 10:1 ratio byweight and allowing the mixture to cure for 7 days attemperatures ranging between 25°C and 35°C. Four mix-tures, comprised of varying concentrations (0, 1, 3 and5% by weight) of 50-�m cellulose particles (Sigmacell�S-5504, Sigma-Aldrich, Oakville, ON, Canada) added asscatterers (Rickey et al. 1995), were tested for pertinentproperties, including the acoustic attenuation, speed ofsound, Young’s elastic modulus and visual appearance,with US B-mode imaging.

The TMM was a variation of the TMM developedby several groups as part of a European Commissionfunded project (Ramnarine et al. 2001; Teirlinck et al.1998; Teirlinck et al. 1997) that used Al2O3 and SiC toachieve suitable attenuation and backscatter properties.For the TMM used here, the agar was replaced with ahigh-gel strength agar as used by Rickey et al. (1995),and formaldehyde was added as a preservative (Blech-inger et al. 1988; Smith et al. 1999). The compositionwas as follows (by % weight): 82.25% distilled water,11.11% glycerol, 2.96% high-gel strength agar (A-6924Sigma-Aldrich), 1.34% formaldehyde (F-1268, Sigma-Aldrich), 0.93% Al2O3 (3 �m, 0CON-008 Calcined,Logitech, Westlake, Ohio), 0.87% Al2O3 (0.3 �m,0CON-015 Ultra Fine, Logitech), and 0.53% SiC (400grit, 17 �m mean, 0CON-003, Logitech).

All of the ingredients, except the formaldehyde,

Thin-walled carotid vessel phantom ● T. L. POEPPING et al. 1069

were mixed together and cooked in a microwave oven at98°C, using the built-in constant-temperature probe, for1 hour. The mixture was partially covered while cooking,to reduce evaporation, and occasionally stirred. The mix-ture was allowed to cool to 48°C before adding theformaldehyde and further cooled to 45°C before casting.

Attenuation and speed of sound measurementsThe acoustic attenuation coefficients of the TMM

and the four Sylgard� mixtures were determined using apulse-transmission technique (See Appendix 1) (Dunnand Goss 1986; Madsen et al. 1982; Rickey et al. 1995;Wells 1977). An US pulse of frequencies ranging from 2to 7.5 MHz and approximately 5-�s duration was pro-duced using a single-element unfocussed transducer anda waveform generator (Wavetek 148A) gated with apulse generator (PG501, Tektronix, Beaverton, OR). Thepulse was transmitted through a sample of known thick-ness in a distilled water bath, both at room temperature(22 � 1°C) before measurements, and received by asecond transducer at a fixed distance in the Fraunhoferzone of the US field. Samples were kept in the water bathfor a minimal time, typically less than 1 min, to avoidany changes due to immersion. The peak-to-peak ampli-tude of the received signal was measured with an oscil-loscope (PM3365A, Philips, Eindhoven, The Nether-lands), after checking that there was no observable dis-tortion due to nonlinear propagation. The pulse durationwas selected so that it was long enough to ensure thattransient pulses were not involved in the peak-to-peakamplitude measurements, but short enough so that pulsesdid not overlap. To investigate the materials of interest,we studied right-angle cylindrical samples ranging inthickness from 0.5 to 40 mm. The face of each samplewas adjusted to be perpendicular to the beam axis byfinding the maximum signal amplitude. For the thinnersamples, this was achieved using a gimbal jig (Selfridge1985), which held the sample flat in the vertical planewith two degrees of rotational freedom in motion for fineadjustment. The logarithm of the peak-to-peak amplitudewas plotted vs. the sample thickness, and the attenuationof the sample (in dB cm�1) was determined from theslope of the line of best fit found using linear regression(Matlab� 6.0, The MathWorks, Inc., Natick, MA). Forthe Sylgard� samples, the attenuation coefficient andassociated error were derived from the slope and itscorresponding error. However, for the TMM, the error inthe attenuation coefficient reflects the standard deviationfound from testing three different batches.

The attenuation coefficients were fitted to the fol-lowing power function of frequency:

b
�� f � � a � f (1)
where � (dB cm�1) is the attenuation coefficient, a (dBcm�1 MHz�b) and b are fitting parameters and f (MHz)is frequency. The fit parameters were derived from theintercept and slope of a linear regression fit to the log-logplot of the attenuation coefficient vs. frequency.

The SoS in the four Sylgard� mixtures and theTMM were determined by measuring the time for a5-MHz pulse of 0.5-�s duration to travel a fixed distance(7 cm) through a sample of known thickness in a 22 �1°C distilled water bath. The transit time was measuredfor eight sample thicknesses ranging from 0.5 to 70 mmand plotted as a function of the sample thickness. TheSoS and associated error were determined from the slopeand the corresponding error in the slope. Additionally, toinvestigate the effect of water exposure, the SoS in theTMM was measured within 2 h after it had solidified andthen monitored over time at intervals of 1, 4 and 24 hafter being immersed in water.

Elasticity measurementsThe elasticity of each of the four Sylgard mixtures

was tested using a material test system (MTS) (858Bionix test system and 448 series controller, MTS Sys-tems Corp., Eden Prairie, MN). Strips were cast of eachmixture in the following configuration: 5 mm wide,approximately 2 mm thick and 15 to 20 mm long. Thethickness (� 0.005 mm) was determined using an elec-tronic caliper, and the zero-stress length (i.e., test lengthbetween sample attachment points with no stress applied,ranging between 4.2 to 11.5 mm) was recorded from theMTS unit. A total of 10 to 16 samples of each mixturewere tested using a preconditioning routine of at leastfive cycles of loading and unloading, at a constant rate ofelongation of approximately 2 mm s�1, with a ramp timeof 1 s. The stress (�) was fitted to a linear-plus-exponen-tial function of strain (�) of the form:

� � A � B� � Cek� (2)

This type of function provides excellent flexibilityfor fitting stress-strain data over a wide range of mate-rials and strain. The average fit was determined from 8 to13 samples of each mixture, as available. From eqn (2),it was also possible to determine Young’s elastic mod-ulus, E, which is defined as the first derivative of thestress-strain curve with respect to strain.

Phantom fabricationThe realistic vessel geometries used here were

based on the anthropomorphic carotid bifurcation modelsderived by Smith et al. (1996). Figure 1 shows an ideal-ized healthy, or normal, geometry and demonstrates a
concentrically formed plaque (grey region) that can build


up in the bulb of the internal carotid artery (ICA), result-ing in a stenosis, in this case, a 60% stenosis by the NorthAmerican Symptomatic Carotid Endarterectomy Trial(NASCET) criteria. Additional models with varying de-grees of stenosis severity or other arbitrary geometriescan be generated. A lost-material casting technique (Po-epping et al. 2002; Smith et al. 1999) was used tofabricate the phantoms as follows. A numerically con-trolled router (Millenium, AXYZ Automation Inc., Bur-lington, ON, Canada) was used to fabricate two concen-tric molds of the desired carotid geometry. The first wasan aluminum mold (which represents the lumen bound-ary) milled with a life-sized model geometry, having acommon carotic artery (CCA) diameter of 8 mm. Thesecond was an acrylic mold milled with a dilated modelgeometry that uniformly extends the life-sized geometryradially by a specified amount to produce the desiredwall thickness, in this case, 1 mm, at all points along thecentral axis of each vessel branch. A lumen model wascast, using a low-melting-point (47.2°C) alloy (Cer-rolow� 117, Cerro Metal Products, Bellafonte, PA) in thealuminum mold. The metal cast was centered within thedilated-geometry acrylic mold, and the Sylgard� mixturewas injected to uniformly fill the 1-mm gap between theacrylic mold and the metal core. After the Sylgard� hadcured, the Sylgard� vessel with metal core was removedfrom the outer mold, mounted in an acrylic box andattached to fixed plastic connectors (APC series, ColderProducts Co., St. Paul, MN). The entire phantom wassealed with an o-ring and lid, consisting of a thin (0.25mm) Lexan� sheet attached to a Lexan� frame. Althoughthe low acoustic velocity of the Lexan� sheet can lead toreverberations in the B-mode image when the US beamis perpendicular to the surface, the signal loss is lower

Fig. 1. Carotid bifurcation models with normal geometry (top)and with a 60% stenosis (bottom) by NASCET criteria, show-ing the CCA branching into the ICA and ECA. The grey regiondemonstrates how the 60% stenosis can be formed from aconcentric build-up of plaque within the bulb of the normal

geometry.

relative to other possible materials because it is available

as a very thin, yet strong, sheet. The box was filledthrough a filling port with the agar-based TMM andallowed to set for 24 h before the metal core was meltedout by submerging the phantom in a 52°C water bath.Only the inner surface of the VMM and metal core, notthe TMM, were directly exposed to the water to avoidany changes in the TMM.

Additionally, this technique allows the simulationof plaques on the vessel wall. The plaques can be fabri-cated by injecting additional elastomer into the bulb ofthe ICA, as depicted by the grey region in Fig. 1. Theplaque echogenicity can be modified by varying thescatterer concentration, relative to the TMM and VMM,to produce hypoechoic or hyperechoic plaques. To dem-onstrate this, a pre-made stenosed-lumen vessel (i.e., a1-mm thick, cured Sylgard� vessel coating the metal castof a stenosed model) was mounted within the dilatedacrylic outer mold of a normal model. The bulb wasfilled with the desired Sylgard� mixture, so that the gapwas filled between the precoated metal core (innerboundary) and the dilated normal-geometry acrylic mold(outer boundary).

Comparison of velocity profilesIn addition to the thin-walled phantom, a wall-less

phantom was fabricated using the same lost-materialcasting technique (Poepping et al. 2002; Smith et al.1999). To perform a comparison of velocity profileswithin the two types of phantoms, the phantoms wereperfused with a blood-mimicking fluid (Ramnarine et al.1998) at a constant flow rate of 5 mL/s. Doppler quadra-ture data were collected in a region of fully developedlaminar flow in the CCA using a semiautomated acqui-sition system (Poepping et al. 2002) (Ultramark 8 withAccess 10-PV transducer, Philips-ATL, Seattle, WA). Ateach site, 1 s of data were collected within a 45 � 45point grid, with 0.2-mm increments, which was trans-verse to the direction of flow. Data were acquired usinga 7000-Hz pulse-repetition frequency, 50-Hz wall filter,5-MHz transducer frequency, 2.1-cm focal point,1.5-mm (smallest available) sample volume setting, and60° Doppler angle. The data were analyzed using a1024-point fast Fourier transform (FFT) with 1024-pointHanning window and 50% overlap between consecutivewindows. The peak and mean velocities were determinedat each site from the average power spectrum from 1 s ofdata (i.e., 85 FFT windows). The mean velocity (vmean) isdetermined from the first moment of the average powerspectrum, and the peak velocity (vpeak) corresponds tothe velocity marking 95% of the total integrated powerfrom the average power spectrum. The predicted para-bolic velocity profile was also calculated based on the
assumption of fully developed laminar flow, for which


the maximum centerline velocity is twice the averagevelocity determined from the flow rate.

RESULTS

Material characterizationThe attenuation coefficient as a function of fre-

quency is shown in Fig. 2 for the four Sylgard� mixtureswith different concentrations of Sigmacell� (0, 1, 3 and5% by weight). The attenuation coefficient varies lin-early with frequency for the pure Sylgard�, but thischanges to a power of approximately 1.25, parameter bin eqn (1), for the Sylgard�-Sigmacell� mixtures, as canbe seen from the fit parameters given in Table 1. Thecorrelation coefficient, r, of the linear regression fit to thelog-log plot of the attenuation coefficient vs. frequency isalso given.

Similarly, Fig. 3 shows the attenuation coefficientof the TMM as a function of frequency. The fit param-eters, given in Table 1, show a nearly linear frequency-dependence resulting in an attenuation rate of approxi-mately 0.56 � 0.03 dB cm�1 MHz�1 at 5 MHz, whichagrees with the rate (0.52 to 0.53 dB cm�1 MHz�1)given by Teirlinck et al. (1998), but is slightly higher

Fig. 2. Attenuation coefficient as a function of frequency forfour VMM mixtures with different concentrations of Sigma-

cell� (0, 1, 3 and 5% by weight, as labeled) in Sylgard�.

Table 1. Acoustic properties for the four vessel-mimmaterial (TMM). The fit parameters, a and b, and th

describes the attenuation coefficient as a function of frefrom the error in the intercept and slope, respectively,

The speed of sound, v, is given, along with the errosample thickness for each VMM and from th

Sample material a (dB cm�1 MHz�b)

VMM with no Sigmacell� 3.5 � 0.3VMM with 1% Sigmacell� 3.1 � 0.1VMM with 3% Sigmacell� 3.8 � 0.3VMM with 5% Sigmacell� 5.7 � 0.3TMM 0.64 � 0.05

than that (0.50 � 0.03 dB cm�1 MHz�1) given byRamnarine et al. (2001). Conversely, the measured speedof sound (1539 � 4 m s�1) for the TMM, also shown inTable 1, agrees with the value (1541 � 3 m s�1) given byRamnarine and colleagues, but is lower than that (1551m s�1) quoted by Teirlinck et al. (1998).

Figure 4 shows the SoS measured in the TMM as afunction of the time immersed in water, demonstratingsignificant differences after 4 h, with a speed of sound(1496 � 3 m s�1) approaching that of water (1487 m s�1

at 22°C) after just 24 h. This is likely due to the diffusionof the glycerol, which primarily determines the SoS, outof the TMM. It should be noted, however, that becausethe differences were not significant at 1 h, it is not likelythat diffusion was rapid enough to result in any signifi-cant error in the initial (t � 0) acoustic measurements.

Table 1 also shows the SoS measured in the fourSylgard� mixtures. The SoS measured in the pure Syl-gard� agrees well with the speed (1027 m s�1) found bySelfridge (1985). Measurements in the samples withgreater concentrations of Sigmacell� were less precisedue to the high attenuation through thick samples ofthese mixtures.

The stress-strain relationship of the Sylgard� isshown in Fig. 5a for the four mixtures. For each mixture,

material (VMM) mixtures and the tissue-mimickingelation coefficient, R, are given for Eqn. (1), whichy for each material. The error in a and b were derivedlog-log plot of attenuation coefficient vs. frequency.

ed from the error in the slope of transit time versusdard deviation of four batches for the TMM

b r v (m s�1)

1.02 � 0.05 0.996 1020 � 201.23 � 0.03 0.999 1030 � 301.31 � 0.06 0.997 1000 � 501.22 � 0.03 0.999 1000 � 500.92 � 0.05 0.996 1539 � 4

Fig. 3. Attenuation coefficient as a function of frequency for theTMM.

ickinge corrquencof the

r derive stan


an elastic hysteresis can be observed with the stressslightly greater during loading (upper curve) than duringunloading (lower curve). This is an indication that not allof the energy is immediately recovered during the un-loading phase. However, it was observed that the elas-tomer slowly relaxes and returns to its original length.The corresponding Young’s elastic modulus, found fromthe slope of the loading curve, as defined in the Methodssection, is shown for each mixture in Fig. 5b. The elasticmodulus ranged between 1.3 and 2.3 MPa overall. How-

Fig. 4. Plot of the speed of sound of the TMM, as a function oftime immersed in water, demonstrating the significant decreaseover time, which is likely due to the diffusion of glycerol out ofthe TMM. The double asterisk indicates a highly significant (p

� 0.001) difference relative to the original sample.

Fig. 5. (a) Stress-strain curves for four VMM mixtures, withdifferent concentrations of Sigmacell� (0, 1, 3 and 5% byweight, as labeled) in Sylgard�, and (b) the elastic modulusderived from the loading (upper) stress-strain curve for each

mixture.

ever, this was limited to a maximum of 1.7 MPa forstrains up to 10%. For comparison, a strain of approxi-mately 6% was observed with an intraluminal pressure of140 mmHg (18.7 kPa) using M-mode US imaging of a1-mm thick, 8-mm diameter Sylgard� vessel.

Figure 6 shows 1-mm thick Sylgard� disks thathave been embedded at a depth of approximately 1.25cm in TMM. The Sylgard� disks were made with Sigma-cell� concentrations varying from 0 to 5% by weight aslabeled. The images were taken with a linear array trans-ducer (ATL L12-5 on HDI 5000 system) positioned atdifferent angles, approximately 80° and 70° for the topand bottom row, respectively, relative to the materialsurface. With nearly perpendicular incidence, interfacesare emphasized due to the increased reflection. However,at angles less than about 75°, the interfaces becomenearly invisible and the materials can only be distin-guished by the difference in backscatter due the differ-ence in scatterer concentration. Hence, the Sigmacell�concentration can be altered to match (or intentionallydiffer from) the scatterer concentration of the TMM forvaried contrast. This is illustrated in Fig. 7, which showsa transverse slice of a Sylgard� vessel embedded in twodifferent tissue-mimicking materials. In Fig. 7a, a sili-cone vessel with 3% Sigmacell� is shown embedded ina TMM previously described by Rickey et al. (1995),which also uses 3% Sigmacell� by weight as a scatteringagent. Figure 7b shows a vessel with 0.5% Sigmacell�embedded in the TMM given in the Methods section. Inboth examples, the image was taken with the transducerat approximately 70° to 80° relative to the phantomsurface and top of the vessel. A lower concentration ofSigmacell� scatterer is desirable with respect to thelower attenuation, as shown by Fig. 2. Shadowing, suchas can be observed in Fig. 7, could be reduced bydecreasing the thickness of the wall.

Fabricated carotid bifurcation flow phantomsTo demonstrate the viability of the silicone-walled

vessel phantom, two stenosed vessel models were fabri-cated and tested. Figure 8 displays a thin-walled, carotidbifurcation vessel cast from Sylgard� mounted in theacrylic box before the addition of the TMM (Fig. 8a) andthe final product sealed with the Lexan� lid (Fig. 8b).Figure 9 is a longitudinal slice obtained by digitallyslicing through a 3-D US B-mode image volume ob-tained using a mechanical linear-scanning system (L3Disoftware, Life Imaging Systems, London, ON, Canada;L12-5, 50-mm linear-array with HDI 5000 machine,Philips-ATL, Seattle, WA).

Figure 10 shows a longitudinal slice from a 3-D USimage volume (B-mode with color-encoded Doppler powerdata) of the second carotid bifurcation phantom, with a
hypoechoic plaque filling the bulb of the normal model to

° relat


surround a 70% concentric stenosis. The hypoechoic plaquewithin the bulb contains Sylgard� elastomer with 0.1% byweight Sigmacell� scattering particles, and the TMM andinnermost vessel wall contain a greater scattering particleconcentration (� 0.5% by weight).

The phantoms showed excellent physical stabilityover time with no gross changes observed over a 2-year

Fig. 6. B-mode images of 1-mm thick strips of seven VMweight, as labeled), embedded in TMM with transdu

transducer angled at about 70

Fig. 7. B-mode image from a transverse slice in a flow phantomwith a 1-mm thick Sylgard�-Sigmacell� vessel, embedded in(a) the TMM previously described by Rickey et al. (1995) with3% by weight of Sigmacell� and (b) the TMM by Teirlinck etal. (1997) with 0.5% by weight of SiC. In both cases, theconcentration of Sigmacell� in the VMM was selected to match

the particle concentration in the TMM as given above.

period. A well-sealed phantom could be left for a fewdays with the lumen exposed to air, or in a water-glycerolbath for long-term storage, without observable changes.This is consistent with the Sylgard� 184 specifications(Dow Corning Corporation 1991) of only 0.1% waterabsorption after 7 days immersion, which is in starkcontrast to the gross changes observed with latex rubbertubing. Unlike previous observations with latex rubber,

ixtures, with different Sigmacell� concentrations (by %sitioned perpendicular to surface (top row) and withive to surface (bottom row).

Fig. 8. Pictures of (a) a Sylgard� carotid vessel mounted in theacrylic phantom box with the metal core prematurely meltedout for demonstration purposes, and (b) the completed phantom

M mcer po

sealed with a Lexan� lid and frame.


no swelling, bulging or sagging was observed in thesilicone vessels.

The phantoms also proved to be mechanically ro-bust. Unlike similar wall-less flow phantoms, the thin-walled vessel phantoms were able to withstand long-termpulsatile mechanical stresses associated with normalphysiological flow waveforms. The thin-walled vesselphantoms were typically used with a realistic carotidflow waveform (Holdsworth et al. 1999) with peak flowof 24 mL/s (6 mL/s mean), but were tested with peakflow up to 35 mL/s (9 mL/s mean). During standardusage, the phantoms were perfused continuously for upto 15 h, and withstood approximately 500 h of physio-logical flow over a 24-month period.

Comparison of velocity profilesThe CCA lumen diameters in the thin-walled and

the wall-less phantoms were measured from multipleM-mode images and found to be 7.96 � 0.03 mm and7.67 � 0.02 mm (mean � SEM), respectively. Due tothe difference in diameters and, hence, different maxi-mum centerline velocities, it is necessary to plot relativevelocity vs. relative radius (i.e., r/R for radial distance rand lumen radius R) to make a valid comparison of thevelocity profiles from the two phantoms. Figure 11shows the horizontal and vertical relative velocity pro-files crossing through the center point of the CCA lumenin the wall-less and thin-walled silicone phantoms, com-

Fig. 9. Longitudinal slice through a 3-D B-mode US volume ofa carotid bifurcation phantom with a 60% stenosis.

pared, in each case, with an ideal parabolic profile. For

each phantom, the velocities shown are relative veloci-ties with respect to the maximum vpeak (denoted as vmax)measured in each phantom (i.e., all values have beenscaled relative to the vmax measured in the correspondingphantom). Good agreement can be seen between thevelocity profiles from the two phantom types. Figure 12shows the point-by-point differences between each mea-sured (vpeak or vmean) profile and a parabolic profilescaled to the corresponding maximum velocity vmax.

The top two plots in Fig. 11a and b show the horizon-tal and vertical relative velocity vpeak profiles (vpeak(r/R)/vmax), which are an excellent match to the parabolic profilenear the center of the lumen. However, near the wall, vpeak

overestimates the expected parabolic velocity profile by upto 7.2 cm s�1 and 8.2 cm s�1 in the wall-less andthin-walled phantoms, respectively, as shown by thepoint-by-point differences in Fig. 12a and b. Similarly,the bottom two plots in Fig. 11c and d show the hori-zontal and vertical relative velocity vmean profiles(vmean(r/R)/vmax), and Fig. 12c and d show the corre-sponding differences between the scaled parabolic pro-file and each of the vmean profiles. The measured vmean

consistently underestimates the parabolic velocity profileby approximately 2 cm s�1 near the center of the lumen,but then overestimates the velocities near the wall by up

�1 �1

Fig. 10. Longitudinal slice through a 3-D US volume (usingpower Doppler with B-mode imaging) from a 70%-stenosed ca-rotid bifurcation phantom with simulated hypoechoic plaque sur-rounding the stenosis. Use of power Doppler to illustrate the flowhelps to demarcate the lumen within the hypoechoic plaque. How-ever, the color does not completely fill the true lumen due to lowvelocity flow resulting in power drop-out along the walls of theCCA, ECA and downstream ICA, which causes the walls toartificially appear thicker. A region of power drop-out can also be

seen downstream of the stenosis due to recirculation.

to 4.2 cm s and 4.9 cm s in the wall-less and


thin-walled phantoms, respectively. With both phantoms,these effects can be attributed to known factors, such asthe finite Doppler sample volume (Bastos et al. 2000),the wall filter, and intrinsic spectral broadening.

Fig. 11. A comparison of the (a) horizontal and (b) verticfrom a wall-less phantom and phantom with an embeddcorresponding to fully developed laminar flow is showphantoms with slightly different lumen diameters, the ve

velocity (vmax) measured in each phantom and plo

Fig. 12. Point-by-point differences between a parabolicprofiles crossing through the center of the wall-less and ton the left and vertical on the right. The absolute velocityprofile and a parabolic profile scaled to the relevant

phantom

DISCUSSION AND SUMMARY

Wall-less phantoms have been the preferred choicefor fabricating flow phantoms with complex or arbitrary

profiles and (c) horizontal and (d) vertical vmean profilesn-walled silicone vessel. The parabolic velocity profilecomparison. For ease of comparison of profiles fromprofiles have been scaled relative to the maximum peaka function of the relative radius (�1 � r/R � 1).

e and measured vpeak (a) and (b) and vmean (c) and (d)lled vessel phantoms. The horizontal profiles are showns represent differences calculated between the measuredum velocity vmax (as measured in the corresponding

al vpeak

ed thin forlocitytted as

profilhin-wa

valuemaxim

).


geometries, such as stenosed lumens or bifurcations.Wall-less phantoms provide an excellent match to softtissue with respect to SoS and acoustic attenuation. How-ever, wall-less phantoms can suffer from rupture andleakage problems under physiological flow conditions,including higher flow rates and intraluminal pressures,especially with highly pulsatile waveforms. Addition-ally, the changes in the SoS of the TMM over time whenimmersed in water, as observed here, indicates that wall-less vessels should be used with care. The effusion effectobserved likely could be minimized by filling and im-mersing the phantom in a 10% glycerol solution forstorage. An effective preservative (such as CA-24 pre-servative, Biochema Schwaben, Memmingen, Germany;or formaldehyde) would also be recommended to preventmold growth for long-term stability.

The thin-walled elastomer phantoms demonstratedhere overcome several of the problems of wall-less phan-toms, along with the limitations of prefabricated tubing.Most importantly, the Sylgard�-based VMM can easilybe cast to create vessels with realistic and arbitrarygeometry, as previously achieved using wall-less vessels(Poepping et al. 2002; Smith et al. 1999), but with thestrength necessary to withstand prolonged physiologi-cally realistic mechanical stresses, unlike wall-less ves-sels. Additionally, the use of a silicone elastomer allowsthe TMM to be sealed from exposure to air or water,unlike wall-less vessels; thus, producing a more stablephantom. Furthermore, a unique feature of the produc-tion technique described here is the ability to simulateatherosclerotic plaques with variable echogenicity rela-tive to the VMM and TMM.

A thin Lexan� sheet was chosen to seal the topwindow of the phantom. Although the high acousticimpedance for Lexan� can lead to reverberations in theB-mode image when the US beam is perpendicular to thesurface, the signal loss is lower relative to other possiblematerials because it is available as a very thin (0.25 mm),yet strong, sheet. The frame fastening the thin sheet tothe phantom is approximately 12.7 mm thick, providingan open well that can be filled with a fluid for acousticcoupling. To prevent refraction errors, a water-glycerolsolution, with acoustic velocity of 1540 m/s, is recom-mended as an acoustic couplant.

The Young’s elastic moduli measured for the Syl-gard� mixtures are well matched to human arterial elas-ticity values, as given in the Introduction section. For theregion of typical in vivo strain values (i.e., � 10%), theYoung’s elastic modulus was linear and appropriate inmagnitude (1.3 to 1.7 MPa) to simulate human arterialelasticity at moderate intraluminal pressures (100 to 130mmHg). The elasticity of the Sylgard� also appears to benonlinear, with an increase at high strains ( 20%).
However, these higher strains are well above those re-
sulting from typical lumen pressures of � 200 mmHg or26 kPa and, hence, are not relevant. Additionally, theTMM has an appropriate compliance, as reported previ-ously (Poepping et al. 2002), with a 6% relative disten-sion, which is representative of the vessel compliance fora 60-year-old human (Hansen et al. 1995).

Compared with soft tissue, the silicone elastomerused here suffers from a low SoS (1020 m s�1 vs.soft-tissue average of 1540 m s�1) and high acousticattenuation (18.1 dB cm�1 vs. 3.0 dB cm�1, at 5 MHz).However, the attenuation is less than that reported forother common vessel-mimicking materials (Hoskins1994), such as C-flex� or latex. Additionally, the atten-uation varies nearly linearly with frequency, unlike C-flex�, which varies with frequency to the power of 1.8(Ramnarine et al. 2001). The IEC 61685 standard by theInternational Electrotechnical Commission (IEC 2001)specifies that the tube wall in an US flow test objectshould reduce the signal by an amount less than thesignal reduction caused by a layer of TMM having athickness equal to the vessel inner diameter (i.d.). The0.5%-Sigmacell� VMM described here, with an attenu-ation of approximately 21 dB cm�1 at 5 MHz, results ina 2.1-dB loss through a 1-mm wall compared to a 2.2-dBloss through 8-mm (i.e., equivalent to i.d. of CCA) ofTMM. However, the smaller i.d. of the ICA and ECAbranches would require a thinner VMM wall to satisfythis IEC standard.

Although the silicone wall may introduce small er-rors due to refraction of the US beam, the good agree-ment between the data from the wall-less and thin-walledphantoms suggests that these errors are negligible overthe majority of the lumen, compared with known errorsdue to the finite Doppler sample volume, wall filter andintrinsic spectral broadening. For example, with a samplevolume centered at the wall, where the velocity is ex-pected to be zero, half of the sample volume extends intothe lumen, thereby including nonzero velocities and,thus, overestimating the velocity at the wall. Also, nearthe wall, the 50-Hz wall filter contributes to furtheroverestimation of the velocities because it acts to sup-press low velocities (� 1.3 cm s�1 for the present Dopp-ler angle and analysis parameters); thus, again increasingthe mean of the velocity distribution. This suggests thatDoppler US velocity measurements very near the wall-lumen boundary must, in general, be interpreted withcaution.

The search for a suitable VMM has proven to bedifficult, with few suitable materials that match all of thecriteria described earlier. One of the primary consider-ations here is the need for a material that allows forcustomized fabrication of a robust flow phantom, witharbitrary vessel wall thickness and geometry, that can
withstand the mechanical stresses associated with phys-


iological flow. Ultimately, a compromise between thedesired acoustic properties and the constraints of practi-cal implementation appears to be unavoidable. For stan-dard US applications, the advantages of the silicone-walled phantoms, described here, appear to outweigh thelimitations.

Future workFor the carotid phantoms shown here, 0.5% Sigma-

cell� was added to the Sylgard� to approximately matchthe concentration of scattering particles in the TMM(0.5% SiC). This results in a VMM with similar back-scatter to the TMM, as seen on the B-mode US images,such as shown in Fig. 6. The Sigmacell� was chosen asa scattering agent because it was found to mix well withthe Sylgard� without signs of clumping. Other scatterers,such as silicon carbide, amorphous silica, aluminum ox-ide or nylon (e.g., Orgasol�), may also work and may beinvestigated in the future.

Although no macroscopic changes were observed inthe materials or lumen geometry, microscopic changesmay have occurred. For applications requiring highlyaccurate geometry, it would be necessary to test thelong-term accuracy of the geometry. However, thiswould require a detailed study with high-resolution im-ages and a very precise segmentation tool, which arebeyond the scope of the current paper.

Finally, the phantom shown here incorporates anidealized carotid bifurcation geometry but, obviously,other arbitrary geometries are possible, including patient-specific, 3-D geometries, which can be derived fromhigh-resolution selective cone-beam angiography sys-tems (Fahrig et al. 1997).

Acknowledgments—The authors acknowledge Andrew Hui for his as-sistance with the material stress testing, Chris Blake and Sue Soney fortheir assistance with the 3-D ultrasound scans, Dr. Paul Picot for hisadvice on acoustic property measurements and ATL (Advanced Tech-nology Laboratories, Philips, Seattle, WA) for the HDI 5000 ultrasoundunit. Financial support has been provided by the Heart and StrokeFoundation of Canada (grant #T-5135 and research scholarship ofD.W. Holdsworth), in conjunction with the Canadian Institutes ofHealth Research (group grant #GR-14973 and CIHR/HSFC partnershipgraduate scholarship of T.L. Poepping), and the Robarts ResearchInstitute (graduate student award of T.L. Poepping). D.W. Holdsworthis a Career Investigator, supported by the Heart and Stroke Foundationof Ontario.

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APPENDIX

In the pulse-transmission technique (Dunn and Goss 1986; Madsen etal. 1982; Rickey et al. 1995; Wells 1977), the emitted signal A0 istransmitted through the water plus sample and received as A1. Theattenuated signal A1 can be written as follows:

A1

A0� T1 · T2�e � �wdwe � �sds�, (A1)

where, respectively, T1 and T2 are the transmission coefficients at thesample entrance and exit surfaces, �s (cm�1) and �w (cm�1) are theamplitude attenuation coefficients of the sample and water, and ds anddw are the distances traveled through the sample and water, so that thetotal distance between the transmitting and receiving transducers isdtotal � (dw ds). The relative change in signal power, in dB, is givenby:

P�dB� � 20 · log10

A0

A1, (A2)

and the attenuation coefficient, in dB cm�1, is given by:

��dB/cm� � 20 � · log10e (A3)

Hence:

log10A1�(�w��s)

20ds(log10T1T2A0�

�w · dtotal

20) , (A4)

and �s can be determined then for each transducer frequency, f (MHz),from the slope of the plot of log of signal amplitude (A1) vs. samplethickness (ds) after determining �w (dB cm�1) using �w � 0.002 f 2

(Dunn et al. 1969; Insana 1995).

A thin-walled carotid vessel phantom for Doppler ultrasound flow studies

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