Noninvasive simultaneous assessment of wall shear rate and wall distension in carotid arteries

Ultrasound in Med. & Biol., Vol. 32, No. 11, pp. 1661–1670, 2006 Copyright © 2006 World Federation for Ultrasound in Medicine & Biology Printed in the USA. All rights reserved 0301-5629/06/$–see front matter

doi:10.1016/j.ultrasmedbio.2006.07.023

● Original Contribution NONINVASIVE SIMULTANEOUS ASSESSMENT OF WALL SHEAR RATE AND WALL DISTENSION IN CAROTID ARTERIES PIERO TORTOLI,* TIZIANO MORGANTI,* GIACOMO BAMBI,* CARLO PALOMBO,† and KUMAR V. RAMNARINE‡ *School of Engineering, University of Florence, Firenze, Italy; †School of Medicine, University of Pisa, Pisa, Italy; and ‡Department of Medical Physics, University Hospitals of Leicester NHS Trust, Leicester, UK (Received 19 December 2005; revised 19 June 2006; in final form 13 July 2006)

Abstract—A novel technique has been developed for the noninvasive real-time simultaneous assessment of both blood velocity profile and wall displacements in human arteries. The novel technique is based on the use of two ultrasound beams, one set at optimal angle for wall motion measurements and the other for blood velocity profile measurements. The technique was implemented on a linear array probe divided into two subapertures. A modified commercial ultrasound machine and a custom PC board based on a high-speed digital signal processor was used to process the quadrature demodulated echo signals and display results in realtime. Flow phantom experiments demonstrated the validity of the technique, providing wall shear rate (WSR) estimates within 10% of the theoretical values. The system was also tested in the common carotid arteries of 16 healthy volunteers (age 30 to 53 y). Results of simultaneous diameter distension and WSR measurements were in agreement with published data. (E-mail: [email protected]) © 2006 World Federation for Ultrasound in Medicine & Biology. Key Words: Multigate Doppler processing, Wall shear rate, Wall distension, Arterial mechanics.

to the mechanical interaction with blood flow is considered to influence the atherogenetic process (GarciaCardena et al. 2001). Site-specific differences in flow distribution inside the vessel, in particular, near the vessel walls, yielding different time-dependent shear stress waveforms, are reported between atherosclerosis-susceptible and atherosclerotic human carotid arteries (Dai et al. 2004). Wall shear stress (WSS), given by the product of the wall shear rate (WSR) and the total blood viscosity, represents the viscous drag exerted on the vessel wall by the flowing blood. The WSS appears to play an important role in acute adaptations to flow changes, vascular remodeling and atherosclerosis. Atherosclerosis has been associated with low mean WSS (Irace 2004; Zhao 2002) and disturbed flow pattern, as well as with large temporal or spatial gradients of WSS in the carotid artery bifurcation (Ku et al. 1985). Determination of flow velocity profile and WSR in superficial large arteries, such as carotid and femoral arteries, may help in identifying sites preferentially prone to atherogenesis. Furthermore, the simultaneous measurement of arterial wall motion and WSS may improve sensitivity and accuracy of assessing endothelial dysfunction (Reneman et al. 2005) by means

INTRODUCTION In developed and emerging economies, cardiovascular disease is a leading cause of death and permanent invalidity (Murray et al. 1997). Working conditions of the vascular system need to be investigated and monitored to help to improve prevention of atherosclerotic related disease as well as detection and treatment of clinical events such as stroke, myocardial ischemia and heart failure. In the usually long preclinical phase of atherosclerosis, the evaluation of arterial properties such as wall thickness, distensibility and stiffness can represent a valuable tool to assess the lesion level or monitor the disease status (O’Rourke et al. 2002). An increased carotid intima-media thickness (IMT) is an established marker of atherosclerotic risk load even in young healthy subjects (Knoflach 2003) and increasing evidence highlights arterial stiffness and endothelial dysfunction as valuable prognostic indicators (Weber et al. 2004; Bonetti and Lerman 2003). The response of endothelium Address correspondence to: Piero Tortoli, Electronics and Telecommunications Dept., University of Florence, Via Santa Marta 3, 50139 Firenze, Italy. E-mail: [email protected] 1661

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of flow-mediated dilation. This approach, although quite popular for clinical use (Moens et al. 2005), is still subject to criticism, due to low spatial resolution and accuracy (De Roos et al. 2003), unless echo-tracking techniques are used (Hijmering et al. 2001). Ultrasound techniques are particularly well suited to the measurement of wall motion and numerous clinical studies have demonstrated the potential clinical value of assessing wall motion in vessels such as the carotid arteries (Hoeks et al. 1990; Brands et al. 1997; Ramnarine et al. 2003; Morganti et al. 2005), femoral artery (Balkestein et al. 2001), brachial artery (Kaiser et al. 2001), cerebral arteries (Hayashi et al. 1980) and aorta (Lehmann 1993). On the other hand, the measurement of WSR still represents a challenging task. Ultrasound techniques offer the potential for high-resolution estimates [e.g., 300 ␮m in the system of Brands et al. (1999)] of displacement and flow velocity, but there are a number of problems due to the tendency of clutter to obscure the low frequency and low amplitude Doppler echoes backscattered from blood flowing near the walls. Indirect WSR estimates, based on the measurement of peak velocity in the vessel center and an assumed velocity profile shape, have been proposed (Gnasso et al. 1996; Lou et al. 1993; Wu et al. 2004). However, such an assumption is frequently unrealistic and more direct methods, based on the measurement of the velocity profile along a section of the vessel (Brands et al. 1999), have to be preferred. Whichever is the selected approach, a major problem is due to the need to measure the actual velocity magnitude and direction, and not just the velocity component in the ultrasound beam direction as provided by classic Doppler measurements. This means that the beam flow angle must also be known. Although a number of vector Doppler techniques have been proposed to estimate this angle by combining the velocity components estimated by two or more transducers (Steel et al. 2003a, 2003b), WSR measurements have typically been based on the manual alignment of the angle cursor with the vessel walls in the B-mode image of the vessel (Samijo et al. 1998; Bambi et al. 2004). Under ideal conditions, subjective alignment to within 2° of the beam-vessel direction is considered to be achievable (Thrush and Hartshorne 2005). However, inadvertent angle drift during acquisition of spectral Doppler data are not uncommon and angle changes of over 20° have been reported under clinically realistic conditions (Steel et al. 2003a). In this paper, a novel dual-beam method coupled to multigate Doppler processing is proposed which not only solves the Doppler angle ambiguity, but also allows the accurate estimation of both WSR and wall distension. The system was applied in vivo in the common carotid

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B

A

v

Fig. 1. Geometry of the multigate dual-beam system.

artery (CCA) and evaluated in vitro in flow phantom experiments. METHODS The novel measurement method is based on the simultaneous use of two transducers, which may be either single element cylindrical probes, or two subapertures in a linear array probe. For each transducer, the corresponding M-line echo data are processed with a multigate approach, by acquiring and processing 128 complex samples corresponding to an equal number of adjacent depths. Let us consider the dual-transducer configuration shown in Fig. 1. The US beam associated with transducer A (hereinafter referred to as the reference beam) is transversely oriented to the vessel of interest. Such an orientation is ideal for an optimal investigation of the wall movements. However, by processing the Doppler shifted echo-signals backscattered from blood flowing inside the vessel, the same transducer can also provide a valuable and accurate estimate of the reference-beam-flow angle. Finally, knowledge of the interbeam angle between the two transducers allows transducer B to perform Doppler measurements in which the real Doppler beam-flow angle is known. By extending the velocity measurement to multiple gates, the velocity profile is obtained and the shear rate inside the vessel can thus be estimated, so that the values close to the near and far walls (WSR) are finally extracted. The following paragraphs describe in detail how the reference beam was employed to measure the vessel diameter and distension as well as the flow direction, while the beam of the other transducer (hereinafter referred to as Doppler beam) was used to estimate blood velocity magnitude and WSR.

Shear rate and distension measurement ● P. TORTOLI et al.

Diameter and distension measurement When transversely oriented to the vessel, the reference beam produces strong echoes in correspondence to the anterior and posterior walls. The approximate positions of the walls were thus roughly identified by finding the depths (gates) at which the A-mode signal presented the maximum gradients. The signal samples selected from the wall gates were processed through a modified 2D-autocorrelation algorithm (Loupas et al. 1995) to extract the instantaneous wall velocities. The integration of such velocities produced the instantaneous displacement of the walls, which were used continuously to update the wall gates positions through a tracking procedure (Hoeks et al. 1990). The average difference between the depths of the wall gates was assumed to be the mean vessel diameter. The numerical value corresponding to the difference between the positive and negative peaks of the diameter variation (distension waveform) was defined as peak-to-peak “distension”. A detailed description of the methods adopted for diameter and distension measurements is given in Morganti et al. (2005). Angle measurement If the reference beam is at a right angle to the flow, the Doppler spectra of backscattered echo-signals are substantially symmetrical around the zero mean frequency. Such a unique property, which was demonstrated through several in vitro and in vivo experiments (Tortoli et al. 1993), is derived from the transducer focusing features, which involve a set of effective beamflow angles equally distributed around the nominal 90° Doppler angle (Censor et al. 1988). The spectral symmetry condition represents a sensitive indicator of the transverse beam-flow orientation. Even small deviations from the ideal right angle cause significant mean frequency shifts and visible losses of spectral symmetry. By simply checking the spectrogram, it is thus possible to control when the desired transverse orientation is actually achieved. Recent experiments have shown that flow direction estimates based on this approach can lead to rms errors of lower than 1° (Tortoli et al. 2005). Once the reference beam has been oriented transversely to the flow direction, the beam-flow angle associated with the Doppler transducer is given by ␪ ⫽ 90° ⫾ ␦, where ␦ is the interbeam angle. Velocity magnitude and WSR measurement The echo signals backscattered from multiple depths along the Doppler beam were processed through the standard 128-point fast Fourier transform (FFT) algorithm. The so-called spectral profile (Tortoli et al. 1997), i.e., a matrix of 128-point power spectral densities

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corresponding to 128 different depths, was thus obtained and displayed in realtime. The availability of a complete spectrum for each investigated sample volume (SV) made it possible to investigate the corresponding velocity, Vmax, through the spectral peak frequency, fmax. In each spectrum, after zeroing the first samples around the zero frequency, fmax was estimated through a modified threshold crossing method (Mo et al. 1988) and then converted to velocity according to the equation: Vmax ⫽

f max 2f 0 cos ␪ ⫹ k sin ␪兲 c 共

,

where f0 is the transmitted frequency, c is the US propagation velocity and k is a factor of proportionality depending on the transducer geometry. The second term in the denominator allows correction for intrinsic spectral broadening effects (Tortoli et al. 1995). For circular aperture transducers, the k factor is well approximated by the ratio between the diameter, W, and the focal length, F. The same equation can also be adapted to the case of steered beams (Hoskins 1999). A polynomial least-square fit was applied on the 128 experimental velocity points obtained in this way, and the resulting profile was used to evaluate the gradient with respect to radius (shear rate). The peak shear rate values, typically obtained from SVs close to the near and far blood-wall interfaces, were taken as the anterior and posterior WSR values, respectively. For pulsatile flow in vivo, the “mean WSR” was calculated as the timeaveraged wall shear rate and the “max WSR” was calculated as the time-averaged wall shear rates at peak systole, over the entire acquisition interval. Experimental set-up Two different test beds were implemented for in vitro and in vivo experiments. In the former case, two single-element pencil probes were employed, while, in the latter, a modified commercial echographic machine (Megas, Esaote SpA, Florence, Italy) was used. The main features of the single element transducers employed in the in vitro experimental tests are reported in Table 1. The software of the US machine was customized to enable the selection of two independent M-lines in the same B-mode display. For each line, the steering angle could be set between ⫺18° and ⫹18° with 6° steps. The maximum relative steering angle was thus 36°, which could correspond, e.g., to beam-flow angles of 90° and 54° for the two lines, respectively. The selection of the two lines involved alternately firing the corresponding subapertures of the 192-element linear-array probe. Outputs from the echograph machine included the two I/Q

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Table 1. Parameters of transducers used in in vitro experiments Transducer

Purpose

Frequency

Bandwidth @-3dB

Aperture

Focus

Beamwidth @-6dB

TXA TXB

Walls, 90° check Flow

8.5 MHz 7.5 MHz

50% 70%

6 mm 6 mm

20 mm 20 mm

0.9 mm 1.2 mm

components of the demodulated echo signal, together with a TX synchronisation pulse. By considering the duration of transmitted pulses (typically 1 to 2 ␮s) and the effect of low-pass filters after demodulation, the overall axial resolution of the Megas system was estimated to be about 1.5 mm. In all experiments, the received echo signals were processed in a digital signal processing (DSP) board, recently developed at the Microelectronic Systems Design Laboratory of the University of Florence (Ricci et al. 2006). During in vitro experiments, the board was used in a stand-alone configuration, i.e., it directly controlled the aforementioned pencil probes in both transmit-receive (TX-RX) modes. For in vivo tests, the board worked only in RX mode, by coherently sampling the echo signals produced by the Megas system (Bambi et al. 2004). In both cases, the received echo-signals were sampled at a rate programmable up to 10 MHz, which corresponds to a range gate spacing of about 75 ␮m. The algorithms described in previous sections were implemented in the real-time acquisition software. The DSP board performed the FFT spectral analysis and the wall distension measurements, while the PC estimated the wall shear rate from the velocity profiles. A more flexible implementation of these algorithms in LabVIEWTM software (National Inst., Austin, TX, USA), with the capability automatically to extract numeric val-

ues of the measured quantities, was also integrated in a postprocessing tool. The system could store the data acquired from both the selected beams in a 32 MByte SDRAM buffer and save these data in two distinct files (one for each direction). The duration of recorded data were dependent on the PRF. For example, at PRF ⫽ 10 kHz, 6 s of data were stored. Flow phantom In vitro measurements were performed using an experimental flow phantom consisting of a 7 mm diameter flow channel through a tissue-mimicking material (TMM) at a depth of 2 cm. The inlet length was approximately 40 cm. A schematic diagram of the experimental set-up is shown in Figure 2. The TMM had suitable acoustic properties (speed of sound 1540 m/s and attenuation 0.5 dB/cm.MHz) and consisted of (%weight): 82.97% water; 11.21% glycerol; 0.46% benzalkoniumchloride; 0.53% 400 grain SiC powder (Logitec Ltd, Glasgow, UK); 0.94% 3 ␮m Al2O3 powder (Logitec); 0.88% 0.3 ␮m Al2O3 powder (Logitec); 3.00% agar (VWR, Dorset, UK). Details of the wall-less flow phantom construction are given elsewhere (Ramnarine et al. 2001). A gear pump (Micropump model 120-000-1100, Concord, CA, USA), driven by a computer-controlled linear amplifier (4020 series, Aerotech, Pittsburgh, PA, USA) was used to generate a steady or pulsatile flow of a blood-mimicking fluid through the flow circuit. The

Florence Multigate Doppler System Pump controller: Steady or Pulsatile

Gear Pump

Wall motion transducer

PC Software and Display

Velocity profile transducer

Wall-less channel Tissue mimicking material Blood mimicking fluid

Water bath Test Tank

Fig. 2. Schematic of experimental flow phantom set-up.

Reservoir

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the flow channel and positioned to receive maximum Doppler signal power by both transducers and a symmetric Doppler spectrum for the transducer perpendicular to the flow. Wall shear rate data were acquired at a range of steady flow rates from 127 to 1050 mL/min, which, assuming a parabolic velocity profile through the 7 mm diameter flow channel, corresponded to peak velocities from 11 to 91 cm/s. The flow rates at five flow pump settings were measured by timed collection of a known volume of BMF. In vivo measurements The in vivo measurements were made on a group of 16 volunteers without cardiovascular diseases (age 30 to 53 y), by using the echographic machine as the front-end. In accordance with the Helsinki declaration, ethical approval for the study was obtained and all subjects gave informed consent to the study. The system was preliminarily tested in the common carotid artery. In each case, the M-line directions were set 1.5 to 2 cm proximal to the bifurcation. The diameter, distension and WSR were measured over at least three beats for each volunteer. By using our custom real-time software, the two US beams were optimally orientated by looking at the features of both spectral profiles. The M-line direction was considered to be at 90° to the flow when the spectral profiles looked symmetrical around the vertical center-line and were bounded by strong clutter echoes (Fig. 3a). The Doppler beam was set at Doppler angles resulting from the difference between 90° and the interbeam angle due to steering. In this case, the spectral profiles assumed typical shapes such as that frozen in Fig. 3b. RESULTS Fig. 3. Instantaneous spectral profiles obtained from the scan lines represented in Fig. 1 during the systolic peak phase. In Fig. 3a, corresponding to the reference (A) line, the clutter corresponding to the arterial walls is indicated by arrows. Fig. 3b corresponds to the Doppler (B) line. The power spectral density is color coded (arbitrary units); depth (mm) and frequency (kHz) are in vertical and horizontal axis, respectively.

blood-mimicking fluid consisted of (%weight): 5 ␮m OrgasolTM (2001UDNAT1 Orgasol, ELF Atochem, Paris, France) particles (1.82%); pure water (83.86%); glycerol (10.06%); Sigma D4876 dextran 185000D (Sigma Ltd, Poole, UK) (3.36%); and ICI synperonic N surfactant (Merck Ltd, Glasgow, UK) (0.9%). Preparation and properties of the blood-mimicking fluid are described in detail elsewhere (Ramnarine et al. 1998, 1999). The two transducer system was clamped above

In vitro experiments A summary of the WSR results for the steady flow experiments, made at an interbeam angle of 35°, is reported in Table 2. The WSR measurements were compared with the WSR values predicted by assuming a parabolic velocity profile in the 7 mm diameter flow

Table 2. Summary of measurements results for the in vitro experiments Parabolic Parabolic flow Measured near Measured far Flow rate flow WSR (1/s) WSR (1/s) (ml/min) Vpeak (cm/s) WSR (1/s) 127 231 400 613 1050

11.0 20.0 34.7 53.1 91.0

63 114 198 304 520

68 ⫾ 3 116 ⫾ 5 228 ⫾ 7 353 ⫾ 10 675 ⫾ 12

⫺61 ⫾ 2 ⫺104 ⫾ 2 ⫺175 ⫾ 3 ⫺237 ⫾ 5 ⫺446 ⫾ 8

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Human study Figures 5 and 6 report the result of the postprocessing carried out on the data acquired for the right CCA of a 39 y old male volunteer. The two M-lines have here been set with a 30° interbeam angle. Figure 5 shows a spectral profile frozen during the systole while the Doppler M-line was set at 60° to the vessel. The maximum frequency points obtained from the spectrum are reported, with the corresponding regression curve superimposed. The shear rate obtained from the velocity profile is also represented, with two horizontal cursors marking positions at which WSRs are estimated. Figure 6 (top traces) shows the displacement and distension waveforms obtained from the acquisition of the M-line oriented at a right angle. The peak-to-peak distension was extracted and averaged over four cardiac cycles, leading to a mean value of 558 ␮m and 22 ␮m standard deviation (SD). The diameter was 6.53 mm and the relative distension was 8.5%. The corresponding anterior and posterior wall shear rate waveforms are shown in the two bottom traces. Over the four cardiac cycles, maximum WSR results were 1169 s⫺1 for the anterior wall and ⫺997 s⫺1 for the posterior wall (with standard deviations within measurement epochs of 74 s⫺1 and 39 s⫺1, respectively). Figure 7 shows the interface of the postprocessing software developed in LabVIEWTM, which allows the simultaneous representation of wall displacement and WSR waveforms, calculated from data acquired along the two lines. Table 3 reports the numerical data estimated for each volunteer. The average distension for the complete

Fig. 4. Spectral profiles obtained during in vitro experiments. Fig. 4a reports an example of asymmetrical profile due to not fully developed flow (flow rate: 1050 mL/min), Fig. 4b reports a typical parabolic profile (flow rate: 127 mL/min).

channel, with peak velocity equal to twice the measured mean velocity. For the two higher flow rate conditions, the flow was not fully developed, due to the insufficient inlet length, so the velocity profile was not parabolic (see Fig. 4a). In the other cases (see, e.g., Fig. 4b), the measured WSR was close to the predicted values, with an overestimation error of less than 10%.

Fig. 5. Spectral profile obtained at peak systole from the right CCA of a 39 y old male volunteer. Doppler angle was 60°. The maximum frequency points obtained from the spectrum are reported, with the regression curve superimposed. The shear rate profile is shown at the right of the spectrum and the two horizontal lines indicate the depths at which the WSRs are estimated.

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Fig. 6. Measurements on the right CCA of the same volunteer of Fig. 5. From top to bottom: displacement of near and far wall and distension waveforms obtained from the M-line set at a 90° beam-vessel angle; WSR of near and far walls obtained from the M-line set at 60° angle.

data set was 464 ␮m (range 261 to 763 ␮m), with a typical 28 ␮m standard deviation within measurement epochs of at least three cardiac cycles each. Both the max (963 versus 824 s⫺1) and mean (335 versus 283 s⫺1) WSR values were higher for the anterior than for the posterior wall. DISCUSSION AND CONCLUSION This paper has presented a novel US system for the simultaneous monitoring of diameter, distension and flow velocity profile at superficial large artery level. In vitro experiments have produced results showing good agreement with predicted WSR values. Preliminary investigation of the common carotid arteries in 16 normal subjects (age: 30 to 53 y) has provided higher maximum and mean WSR values for the anterior wall compared with the posterior wall, confirming a marked asymmetry in the CCA velocity profiles (Tortoli et al. 2003). Previous studies have generally given WSR values as the average of the anterior and posterior WSR (Samijo

et al. 1997). In our case, by averaging the available WSR values over anterior and posterior walls, and over 16 subjects, maximum and mean WSR values of 892 ⫾ 167 s⫺1 and 309 ⫾ 72 s⫺1, respectively, are obtained. Hoeks et al. (1995) reported respective maximum and mean WSR values of 1100 ⫾ 231 s⫺1 and 342 ⫾ 48 s⫺1 in male subjects (age 20 to 30 y) to 765 ⫾ 128 and 310 ⫾ 79 s⫺1 in 60 to 70 y old males. Samijo et al. (1997) reported respective maximum and mean WSR values of 1050 ⫾ 264 s⫺1 and 408 ⫾ 83 s⫺1 in male subjects (age 20 to 29 y) to 767 ⫾ 114 and 357 ⫾ 81 s⫺1 in 50 to 59 y old males and 965 ⫾ 182 s⫺1 and 375 ⫾ 52 s⫺1 in female subjects age 20 to 29 y to 793 ⫾ 136 and 341 ⫾ 99 s⫺1 in 50 to 59 y old females. Kornet et al. (1998) reported respective maximum and mean WSR values of 900 s⫺1 and 310 s⫺1 in 53 subjects age 18 to 67 y. The largest study by Samijo et al. (1998) on 111 subjects age 10 to 59 y reported respective maximum and mean WSR values of 1150 ⫾ 191 s⫺1 and 450 ⫾ 85 s⫺1 in females age 10 to 19 y (1389 ⫾ 219 s⫺1 and 475 ⫾ 68 s⫺1 in males

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Fig. 7. Result of the analysis obtained for a 31 y old female volunteer, presented with the program interface carried out in LabVIEWTM. The upper area is referred to the wall displacement estimation, while the bottom area concerns the WSR estimation, carried out on the data simultaneously acquired over the two selected M-lines.

age 10 to 19 y), decreasing to 793 ⫾ 136 s⫺1 and 341 ⫾ 99 s⫺1 in females 50 to 59 y (767 ⫾ 114 s⫺1 and 357 ⫾ 81 s⫺1 in males 50 to 59 y). Dammers et al. (2003) reported respective maximum and mean WSR values of 1047 ⫾ 345 s⫺1 and 359 ⫾ 111 s⫺1 in 10 subjects age 20 to 30 y. More recently, Wu et al. (2004), who employed a high resolution magnetic resonance imaging (MRI) phase contrast measurement technique, reported respective maximum and mean WSR values of 948 ⫾ 124 s⫺1 and 333 ⫾ 61 s⫺1 in 20 young subjects (mean age 24 ⫾ 2 y).

Our system is characterized by the simultaneous use of two transducers, one aligned perpendicular to the flow and used for estimation of arterial wall distension and diameter and the second transducer set at an inclined angle to the vessel for velocity and WSR estimation based on maximum frequency estimation. As opposed to other triangulation methods, which extract the information on the flow direction from the combination of multiple Doppler measurements, the two transducers play totally different roles here. One is dedicated to the measurement of the Doppler angle, while the other is used for

Table 3. Anterior and posterior wall maximum and mean shear rates, diastolic diameter and distension estimated for a population of 16 volunteers (age 30 –53 y) Subject

WSR (s⫺1) maximum anterior

WSR (s⫺1) maximum posterior

WSR (s⫺1) mean anterior

WSR (s⫺1) mean posterior

Diameter (mm)

Distension (␮m)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 Average values

970 686 1026 973 518 783 983 1146 1169 995 858 1192 981 1315 1040 766 963

⫺913 ⫺975 ⫺797 ⫺1018 ⫺391 ⫺692 ⫺993 ⫺686 ⫺997 ⫺773 ⫺687 ⫺983 ⫺762 ⫺950 ⫺900 ⫺664 ⫺824

228 233 421 340 218 314 366 384 428 323 272 368 323 553 328 264 335

⫺247 ⫺321 ⫺320 ⫺450 ⫺88 ⫺287 ⫺367 ⫺243 ⫺294 ⫺227 ⫺221 ⫺337 ⫺301 ⫺327 ⫺282 ⫺212 ⫺283

7.69 9.16 6.26 6.05 7.07 7.86 6.89 6.90 6.53 7.40 7.16 6.58 6.86 5.87 6.60 6.96 6.99

414 367 581 261 655 368 413 470 558 580 763 426 292 424 521 334 464

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the velocity magnitude estimates. The accurate estimation of the beam-flow angle is achieved by imposing a transverse orientation of the first beam to the flow. Such orientation is easily recognized by inspecting the features of the Doppler spectra which are correspondingly obtained. Transverse alignment with the vessel walls might also be performed by maximizing the echo amplitude from the vessel walls. However, the transverse orientation to the vessel walls, although optimal for distension measurements, is not always coincident with a transverse beam-flow angle. This coincidence is valid only when the flow is parallel to the vessel axis but not, e.g., near stenoses or vessel bifurcations. The technique, whilst primarily designed to provide the ability to estimate velocity and wall distention simultaneously, may, in practice, have advantages over conventional dual-beam vector Doppler techniques for the estimation of angle independent velocity. The implementation means that the possibility for SV distortion and misregistration between beams passing through different overlying tissues may be less likely, as the interbeam angle is effectively halved, whilst the fluctuation in the velocity estimate is not subject to magnification of bias as would be the case in a standard dual-beam vector Doppler implementation (Steel et al. 2003a). The ability of the technique described in this paper to align the beam-vessel angle to better that 1° compares favorably with the fluctuation in beam-vessel angle estimates obtained using dual-beam vector Doppler techniques (Steel et al. 2003a, 2003b). The availability of the complete spectrum for each investigated SV means that it is possible to investigate the corresponding velocity using not only the maximum Doppler frequency as used in this paper, but also the spectral mean and minimum frequencies, which may also be of value. Important considerations include the robustness and accuracy of the estimates to effects such as beam distortion due to inhomogeneous overlying tissue, tissue attenuation, velocity gradient spectral broadening, fluctuation of the velocity estimate and bias due to spatial averaging of the velocity over the finite SV size. Studies have evaluated the use of different Doppler spectral estimates in dual-beam techniques for estimation of angle-independent maximum velocity (Steel et al. 2003a, 2003b). These suggested that, although the maximum frequency estimate is subject to greater velocity fluctuations than the mean frequency estimate, use of the maximum frequency estimate for the estimation of peak flow velocity has advantages over use of the mean frequency estimate, which is more sensitive to SV misregistration and spatial averaging bias (Steel et al. 2003a, 2003b). The possibility simultaneously to monitor WSR and large artery diameter and distension provides new oppor-

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tunities to explore the relation between arterial remodeling and susceptibility to atherosclerosis in patients with risk factors. Jiang et al. (2000), by using a conventional echo-Doppler approach, observed that both low WSR and increased circumferential wall tension were associated with increased arterial diameter and reduced distension in patients with risk factors for atherosclerosis. New perspectives for investigating endothelial function by means of the so-called flow-mediated vasodilation (FMD) are also open (Benjamin et al. 2004; Sidhu et al. 2002). The adequacy of brachial artery vasodilation (i.e., the physiological response) with the entity of stimulus (the change in WSR during reactive hyperemia) can be investigated to assess the efficiency of the biologic signal (release and bioavailability of nitric oxide). In conclusion, this paper has presented a novel ultrasound system capable of simultaneously estimating the WSR and the wall movements in large arteries. The preliminary tests of the system have demonstrated the capability of rapidly obtaining the optimal set-up. The results indicate that the system is a promising tool for clinical studies assessing the relation between the dynamic properties of the vessel wall and the shear rate associated with different flow patterns. Acknowledgments—The authors thank F. Vittone (University of Pisa, Italy) for valuable help in the evaluation of preliminary results. Technical support by F. Andreuccetti (Esaote SpA, Firenze, Italy) is also acknowledged. The authors are grateful for funding by MIUR (COFIN 2005) and by the British Council and the Italian Ministry for Education, Universities and Research (MIUR) in collaboration with the Conferenza dei Rettori delle Universita Italiane (CRUI) Partnership Programme.

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Volume 32, Number 11, 2006 O’Rourke MF, Staessen JA, Vlachopoulos C, Duprez D, Plante GE. Clinical applications of arterial stiffness; definitions and reference values. Am J Hypertens 2002;15:426 – 444. Ramnarine KV, Nassiri DK, Hoskins PR, Lubbers J. Validation of a new blood-mimicking fluid for use in Doppler flow test objects. Ultrasound Med Biol 1998;24:451– 459. Ramnarine KV, Hoskins PR, Routh HF, Davidson F. Doppler backscatter properties of a blood-mimicking fluid for Doppler performance assessment. Ultrasound Med Biol 1999;25(1):105–110. Ramnarine KV, Anderson T, Hoskins PR. Construction and geometric stability of physiological flow rate wall-less stenosis phantoms. Ultrasound Med Biol 2001;27(2):245–250. Ramnarine KV, Hartshorne T, Sensier Y, et al. Tissue Doppler Imaging of carotid plaque wall motion: A pilot study. Cardiovascular Ultrasound 2003;1:17. Reneman RS, Meinders JM, Hoeks APG. Noninvasive ultrasound in arterial wall dynamics in humans: What have we learned and what remains to be solved. Eur Heart J 2005;26:960 –966. Ricci S, Boni E, Guidi F, Morganti T, Tortoli P. A Programmable real-time system for development and test of new ultrasound investigation methods. IEEE Trans Ultrason Ferroelect Freq Contr, in press. Samijo SK, Willigers JM, Brands PJ, Barkhuysen R, Reneman RS, Kitslaar PJEHM, Hoeks APG. Reproducibility of shear rate and shear stress assessment by means of ultrasound in the common carotid artery of young human males and females. Ultrasound Med Biol 1997;23:583–590. Samijo SK, Willigers JM, Barkhuysen R, et al. Wall shear stress in the common carotid artery as function of age and gender. Cardiovasc Res 1998;39:515–522. Sidhu JS, Newey VR, Nassiri DK, Kaski JC. A rapid and reproducible on line automated technique to determine endothelial function. Heart 2002;88(3):289 –292. Steel R, Fish PJ, Ramnarine KV, Criton A, Routh HF, Hoskins PR. Velocity fluctuation reduction in vector Doppler ultrasound using a hybrid single/dual-beam algorithm. IEEE Trans Ultrason Ferroelectr Freq Control 2003a;50(1):89 –93. Steel R, Ramnarine KV, Davidson F, Fish PJ, Hoskins PR. Angleindependent estimation of maximum velocity through stenoses using vector Doppler ultrasound. Ultrasound Med Biol 2003b; 29(4):575–584. Tortoli P, Guidi G, Pignoli P. Transverse Doppler spectral analysis for a correct interpretation of flow sonograms. Ultrasound Med. Biol 1993;19(2):115–121. Tortoli P, Guidi G, Newhouse VL. Improved blood flow estimation using the maximum Doppler frequency. Ultrasound Med Biol 1995;21(2):527–532. Tortoli P, Guidi G, Berti P, Guidi F, Righi D. An FFT-based flow profiler for high-resolution in vivo investigations. Ultrasound Med Biol 1997;23(6):899 –910. Tortoli P, Michelassi V, Bambi G, Guidi F, Righi D. Interaction between secondary velocities, flow pulsation and vessel morphology in the common carotid artery. Ultrasound Med Biol 2003; 29(3):407– 415. Tortoli P, Bambi G, Ricci S. A novel dual beam approach for removing Doppler angle ambiguity. Proceedings of the IEEE Ultrasonics Symposium. Rotterdam 2005:150 –153. Thrush A, Hartshorne T. Peripheral vascular ultrasound, 2nd ed. Churchill Livingstone, 2005. Weber T, Auer J, O’Rourke MF, et al. Arterial stiffness, wave reflections and the risk of coronary artery disease. Circulation 2004;109: 184 –189. Wu SP, Ringgaard S, Oyre S, Hansen MS, Rasmus S, Pedersen EM. Wall shear rates differ between the normal carotid, femoral and brachial arteries: An in vivo MRI study. J Magn Reson Imaging 2004;19:188 –193. Zhao SZ, Ariff B, Long Q, et al. Interindividual variations in wall shear stress and mechanical stress distributions at the carotid artery bifurcation of healthy humans. J Biomech 2002;35:1367–1377.