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(Radiology. 2001;218:109-117.)
© RSNA, 2001

Experimental Studies

Power Doppler US: Evaluation of the Morphology of Stenoses with a Flow Phantom1

Michel Claudon, MD, Daniel Winninger, BSc, Serge Briançon, MD and Patrick Pesque, PhD

1 From the Department of Radiology, Hôpital de Brabois, Rue du Morvan, 54511 Vandoeuvre les Nancy, France (M.C.); the Departments of Technical Assistance (D.W.) and Statistics (S.B.), University of Nancy, France; and the Department of Research and Development, Advanced Technology Laboratories, Bothell, Wash (P.P.). From the 1999 RSNA scientific assembly. Received July 13, 1999; revision requested August 25; final revision received April 27, 2000; accepted May 22. Address correspondence to M.C. (e-mail: michel.claudon@wanadoo.fr).


   ABSTRACT
TOP
 ABSTRACT
INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 
PURPOSE: To determine the importance of technical settings at power Doppler ultrasonography (US) for the evaluation of stenoses.

MATERIALS AND METHODS: A computer-controlled phantom was used to induce a reproducible flow across a calibrated 71% stenosis in an 8.4-mm-diameter tube. Two transducers, 2–4 and 5–10 MHz, working at depths of 3.0 and 11.5 cm, respectively, with different beam angles (40°, 60°, and 90°), were used to simulate evaluation of pulsatile flow across normal and stenotic vessels in various conditions. For each condition, gain, pulse repetition frequency, and wall filter were progressively turned from low to high values. Two observers measured in a blinded fashion the apparent lumen of the stenotic and normal vessels on longitudinal and transverse images with the use of power Doppler US.

RESULTS: When the high-frequency transducer was used, gain significantly affected both stenotic and feeding vessel measurement, whereas pulse repetition frequency and filter only affected feeding vessel evaluation. When the low-frequency transducer was used, all factors, including flow velocity and beam angle, played a significant role (P < .001). In most conditions, overestimation of the lumen and underestimation of the lumen of the feeding vessel led to severe underestimation of the degree of stenosis.

CONCLUSION: Power Doppler US cannot be used to measure stenoses accurately. Underestimation of the degree of the stenosis was significantly higher with the low-frequency probe than with the high-frequency probe.

Index terms: Blood, flow dynamics, 9*.129892 • Blood vessels, stenosis or obstruction, 9*.12989 • Blood vessels, US, 9*.12984, 9*.12989 • Phantoms, 9*.12989 • Ultrasound (US), experimental studies, 9*.12984, 9*.12989 • Ultrasound (US), power Doppler studies, 9*.12989


   INTRODUCTION
TOP
 ABSTRACT
 INTRODUCTION
MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 
Power Doppler ultrasonography (US) has been emerging as a useful adjunct to conventional color Doppler US for the evaluation of vascular diseases and vascularity of organs (1). Instead of color Doppler US, which is based on the mean Doppler frequency shift, power Doppler US displays the integrated power of the Doppler signal, which is primarily related to the number of red blood cells that produce the Doppler shift (2). Power Doppler US has several theoretic advantages over color Doppler US: immunity to aliasing, quasiindependence of the beam angle, and improved depiction of low-velocity flow (1–4).

Clinical application of power Doppler US is being evaluated in various fields, including native kidney and renal transplant diseases, hepatic tumors, testicular torsion, extra- and intracranial arterial disease, and soft-tissue inflammation (1,5,6). The flow information obtained from power Doppler US has also been used for quantitative studies, such as fractional blood flow evaluation (7), tumor vascularity quantification (8), detection of volume flow differences (9), three-dimensional reconstruction studies (10), and enhanced studies (11).

Power Doppler US, in combination with color Doppler US, which displays hemodynamic information, has been promising in the assessment of vascular stenosis: power Doppler US is better for differentiation of surface plaque morphology, and it better depicts low-, middle-, and high-grade stenosis, compared with color Doppler US, in the evaluation of internal carotid arterial (12,13) and renal arterial stenosis (14). However, the correlation between power Doppler US and color Doppler US is moderate for the measurement of the diameter of stenosis (12–14). Overestimation of the vessel lumen with power Doppler US has been described in phantom studies (15) and reported in renal arterial stenosis, although it has not been quantified (14).

As with color Doppler US, power Doppler US is influenced by instrument settings, including pulse repetition frequency and filtering (16). Rubin et al (2) pointed out the role of the optimization of gain. Hemodynamics, insonation angle, and image orientation may also contribute to the results (10,15). However, in most of these experimental studies, the authors did not evaluate stenotic vessels, were focused on superficial stenosis, or studied only a few of the potentially involved factors.

We therefore conducted a power Doppler US study by using a computer-controlled flow phantom to determine the role of various technical factors, including flow velocity, insonation angle, image orientation, and US instrument settings, in the evaluation of the degree of vascular stenosis.


   MATERIALS AND METHODS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 
Experimental System
Figure 1 is a block diagram of our experimental system. The test phantom was composed of a Plexiglas box containing a block of US gel pad (Proxon MAZ2; Kretztechnik, Zipf, Austria), which has acoustic properties similar to those of human tissues—attenuation of 0.35 dB/cm at 4 MHz and US waveform propagation velocity of 1,400 m/sec. On opposite sides of this block, holes were created by using an 8.4-mm copper tube to remove calibrated cylindrical plugs of the US gel pad, while preserving a space of 14 mm between the opposite plugs. A 2.4-mm copper tube was then used to make a hole of a smaller diameter between the two larger holes. This resulted in a 14-mm long stenosis of 71%. The box was filled with water, and its walls were covered with polyurethane absorbing layers (attenuation of 3.5 dB/mm) to minimize reverberation artifacts.



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  		Figure 1. Block diagram of the experimental system. PC = personal computer.

 
The blood-mimicking fluid was a 33% glycerol mixture with starch particles (3–5-µm diameter), which provide the reflecting targets for the ultrasound beam (17). The viscosity of this solution was 4 mPa · sec. The main reservoir containing the liquid was continuously agitated to keep the solution homogeneous. Polyurethane tubes 10 mm in diameter were used to provide an inlet and outlet for the stenosis phantom.

An electronically regulated syringe injector with a 130-mL syringe (Mark IV; Medrad, Indianola, Pa) was modified for the experiment. The motor pump was monitored with a personal computer (T3200; Toshiba, Tokyo, Japan) by using commercially available software (TURBO C 1.5; Borland International, Scotts Valley, Calif) to produce a calibrated pulsatile flow. Each cycle was composed of 10 consecutive phases, all individually programmed on the computer. The displacement of the injector rotating coil was continuously analyzed with a digital counting sensor, and the data were transferred to the computer as a feedback control. Different levels of resistance to the propagation of the fluid were attained by adding some calibrated strictures beyond the stenosis. Compliance was regulated by adding a lateral reservoir at the same level.

The maximum number of consecutive available cycles was the ratio of the syringe volume (130 mL) to the flow volume during each cycle. The first four cycles were used to ensure a stable flow, and the subsequent 7–15 cycles were available to collect data. When the syringe was empty, the liquid was aspirated from the main reservoir back to the syringe through a collateral tube, which was opened by means of an electromagnetic clamp monitored by the computer. This system produced flow with velocities of 0.2–3.0 m/sec in the stenosis and resistive index values of 0.4–1.0. Because the US gel pad is a compliant material, both the channel and the stenosis expanded during systole so that their diameters were approximately 10% larger during systole than they were during diastole.

Methods
We used a US unit (HDI 3000; Advanced Technology Laboratories, Bothell, Wash) with a linear 5–10-MHz and a convex 2–4-MHz transducer. Each transducer was fixed at the arm of a tracing table and submerged into the water of the Plexiglas box. The use of a tracing table allowed precise positioning of the transducer along the stenosis (18).

The US gel pad containing the stenosis was placed within the Plexiglas box at depths of 3.0 and 11.5 cm, respectively, when the 5–10-MHz linear probe and the 2–4-MHz convex transducer were used. Beam focusing was adjusted to the level of the stenosis. In both cases, the computer parameters were set to obtain a flow with a resistive index of 0.6, similar to that usually found in low-resistance organs, such as the internal carotid and renal arteries (Fig 2).



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  		Figure 2. Sagittal power Doppler US image shows spectral waveforms obtained with the high-frequency transducer with the following settings: longitudinal view; insonation angle, 60°; peak systolic velocity, 1 m/sec; resistive index, 0.6; gain, 30%; pulse repetition frequency, 1,500 Hz; filter, low; and color versus echo priority, 100%.

 
The pump settings were adjusted to produce angle-corrected peak systolic velocities of 1 and 2 m/sec in the stenotic vessel. Insonation angles of 40°, 60°, and 90° and longitudinal and transverse views were attained by tilting the US gel pad and adjusting the position of the probe with the tracing table. Since the probe and the stenosis were initially placed in the rotation axis of the tracing table, it was easy to use the 90° rotation capability of the tracing table to move the system from the longitudinal view to the transverse one. This allowed us to obtain each series from both longitudinal and transverse views.

When the high-frequency transducer was used, the baseline instrument settings were as follows: gain, 30%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%.

Series were obtained with the following combinations: longitudinal view, insonation angle of 60°, peak systolic velocity of 1 m/sec; longitudinal view, insonation angle of 60°, peak systolic velocity of 2 m/sec; and transverse view, insonation angle of 60°, peak systolic velocity of 1 m/sec. For each series, the effect of changing one technical setting was analyzed, while keeping the other constant at baseline: gain was progressively increased in a stepwise manner by 10% steps from 20% to 70%; pulse repetition frequency was increased from 500 to 700, 1,000, 1,500, 3,000, and 6,000 Hz; and filters were increased from low to medium, high, and maximum. In addition, to evaluate the effect of the insonation angle, gain was increased by 10% steps from 20% to 70% for longitudinal images with a peak systolic velocity of 2 m/sec for three insonation angles: 40°, 60°, and 90°.

When the low-frequency transducer was used, the baseline instrument settings were as follows: gain, 60%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%. We used a basic gain level of 60% to get a sufficient baseline level of signal to study the changes induced by step variations of pulse repetition frequency and filter. The following series were obtained: longitudinal view, insonation angle of 60°, peak systolic velocity of 1 m/sec; longitudinal view, insonation angle of 60°, peak systolic velocity of 2 m/sec; and transverse view, insonation angle of 60°, peak systolic velocity of 2 m/sec.

On the transverse images, a peak systolic velocity of 2 m/sec was chosen to allow a substantial level of signal at the level of the feeding vessel. For each series, gain was progressively increased by 10% steps from 20% to 80%; pulse repetition frequency was progressively increased from 500 to 700, 1,000, 1,500, and 3,000 Hz; and filters were increased from low to medium, high, and maximum, the other settings being set at baseline. To evaluate the effect of the insonation angle, gain was increased by 10% steps from 20% to 80% for three angles (30°, 60°, and 90°), the other factors being longitudinal view and peak systolic velocity of 2 m/sec.

In all series, the color scale and the level of persistence were kept constant. To minimize measurement error, the calculated zoom was used.

For each measurement, the observers (M.C., D.W.) used the following protocol: (a) Review, frame by frame, a clip corresponding to a series of at least five consecutive cycles. (b) Select the image on which the diameters of the vessel and stenosis were maximal. (c) Measure the diameters of the stenosis, d, and of the feeding vessel, D, by using the hand calipers. The outlines of the stenotic and feeding vessels were defined as the lines of darkest color at power Doppler US. (d) Print the picture and log the results on a personal computer. The degree of the stenosis, S, was then automatically calculated as follows: S = (D - d)/D.

One hundred twenty series of measurements were performed with the high-frequency transducer and 138 with the low-frequency transducer. To evaluate the interobserver variability, all these measurements were performed in a blinded fashion by the two operators. They were repeated after 1 month to evaluate the intraobserver variability.

Statistical Analysis
Comparisons were performed with analysis of variance for repeated measurements to evaluate the intra- and interobserver variability and the effect of each tested parameter. For all tests, a P value less than .05 was considered to indicate a statistically significant difference. If statistical analysis revealed that an experimental condition could not be evaluated reliably, the data of that specific series were subsequently excluded from the overall analysis.


   RESULTS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 
General Analysis
Some examples of changes in patterns of the stenotic feeding vessels and measurements of the diameters, observed on different series, are illustrated in Figures 3–5. The changes of the apparent lumen of normal and stenotic vessels and the degree of stenosis obtained for all the setting variations are represented on curves (Figs 6–8).



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  		Figure 3a. Sagittal power Doppler US images show the effects of stepwise changes in gain on both feeding and stenotic vessels with the high-frequency transducer. Baseline settings are as follows: longitudinal view; insonation angle, 60°; peak systolic velocity, 1 m/sec; pulse repetition frequency, 1,500 Hz; filter, low; and color versus echo priority, 100%. Gain is increased from (a) 20% to (b) 40% and (c) 60%, which results in increasing apparent lumen of the stenotic and normal vessels from 0.30 and 0.65 cm to 0.36 and 0.79 cm and 0.53 and 0.93 cm, respectively.

 


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  		Figure 3b. Sagittal power Doppler US images show the effects of stepwise changes in gain on both feeding and stenotic vessels with the high-frequency transducer. Baseline settings are as follows: longitudinal view; insonation angle, 60°; peak systolic velocity, 1 m/sec; pulse repetition frequency, 1,500 Hz; filter, low; and color versus echo priority, 100%. Gain is increased from (a) 20% to (b) 40% and (c) 60%, which results in increasing apparent lumen of the stenotic and normal vessels from 0.30 and 0.65 cm to 0.36 and 0.79 cm and 0.53 and 0.93 cm, respectively.

 


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  		Figure 3c. Sagittal power Doppler US images show the effects of stepwise changes in gain on both feeding and stenotic vessels with the high-frequency transducer. Baseline settings are as follows: longitudinal view; insonation angle, 60°; peak systolic velocity, 1 m/sec; pulse repetition frequency, 1,500 Hz; filter, low; and color versus echo priority, 100%. Gain is increased from (a) 20% to (b) 40% and (c) 60%, which results in increasing apparent lumen of the stenotic and normal vessels from 0.30 and 0.65 cm to 0.36 and 0.79 cm and 0.53 and 0.93 cm, respectively.

 


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  		Figure 4a. Transverse power Doppler US images show the effects of pulse repetition frequency changes on feeding vessel measurement with the high-frequency transducer. Baseline settings are as follows: transverse view; insonation angle, 40°; peak systolic velocity, 1 m/sec; gain, 30%; filter, low; and color versus echo priority, 100%. Pulse repetition frequency is increased from (a) 500 Hz to (b) 1,000 Hz, (c) 3,000 Hz, and (d) 6,000 Hz, which results in marked decrease in display of flow within the apparent lumen of the normal vessel.

 


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  		Figure 4b. Transverse power Doppler US images show the effects of pulse repetition frequency changes on feeding vessel measurement with the high-frequency transducer. Baseline settings are as follows: transverse view; insonation angle, 40°; peak systolic velocity, 1 m/sec; gain, 30%; filter, low; and color versus echo priority, 100%. Pulse repetition frequency is increased from (a) 500 Hz to (b) 1,000 Hz, (c) 3,000 Hz, and (d) 6,000 Hz, which results in marked decrease in display of flow within the apparent lumen of the normal vessel.

 


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  		Figure 4c. Transverse power Doppler US images show the effects of pulse repetition frequency changes on feeding vessel measurement with the high-frequency transducer. Baseline settings are as follows: transverse view; insonation angle, 40°; peak systolic velocity, 1 m/sec; gain, 30%; filter, low; and color versus echo priority, 100%. Pulse repetition frequency is increased from (a) 500 Hz to (b) 1,000 Hz, (c) 3,000 Hz, and (d) 6,000 Hz, which results in marked decrease in display of flow within the apparent lumen of the normal vessel.

 


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  		Figure 4d. Transverse power Doppler US images show the effects of pulse repetition frequency changes on feeding vessel measurement with the high-frequency transducer. Baseline settings are as follows: transverse view; insonation angle, 40°; peak systolic velocity, 1 m/sec; gain, 30%; filter, low; and color versus echo priority, 100%. Pulse repetition frequency is increased from (a) 500 Hz to (b) 1,000 Hz, (c) 3,000 Hz, and (d) 6,000 Hz, which results in marked decrease in display of flow within the apparent lumen of the normal vessel.

 


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  		Figure 5a. Sagittal power Doppler US images show the effects of filter stepwise changes on both feeding and stenotic vessels with the low-frequency transducer. Baseline settings are as follows: longitudinal view; insonation angle, 60°; peak systolic velocity, 2 m/sec; gain, 60%, pulse repetition frequency, 1,500 Hz; and color versus echo priority, 100%. The filter is increased from (a) low to (b) medium and (c) maximum, which results in considerable reduction in signal in the feeding vessel.

 


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  		Figure 5b. Sagittal power Doppler US images show the effects of filter stepwise changes on both feeding and stenotic vessels with the low-frequency transducer. Baseline settings are as follows: longitudinal view; insonation angle, 60°; peak systolic velocity, 2 m/sec; gain, 60%, pulse repetition frequency, 1,500 Hz; and color versus echo priority, 100%. The filter is increased from (a) low to (b) medium and (c) maximum, which results in considerable reduction in signal in the feeding vessel.

 


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  		Figure 5c. Sagittal power Doppler US images show the effects of filter stepwise changes on both feeding and stenotic vessels with the low-frequency transducer. Baseline settings are as follows: longitudinal view; insonation angle, 60°; peak systolic velocity, 2 m/sec; gain, 60%, pulse repetition frequency, 1,500 Hz; and color versus echo priority, 100%. The filter is increased from (a) low to (b) medium and (c) maximum, which results in considerable reduction in signal in the feeding vessel.

 


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  		Figure 6a. Graphs show results obtained with the high-frequency transducer. Baseline instrument settings are as follows: gain, 30%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%; and insonation angle, 60°. The reference diameters of the feeding vessel (DO = 0.84 cm) and stenotic vessel (do = 0.24 cm) are represented on each graph as dashed and dotted lines, respectively. Variations of measurement of the diameter of the feeding vessel (bold line) and the diameter of the stenotic vessel (lighter line) are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), and (c) filter for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); and transverse view, peak systolic velocity of 1 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|).

 


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  		Figure 6b. Graphs show results obtained with the high-frequency transducer. Baseline instrument settings are as follows: gain, 30%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%; and insonation angle, 60°. The reference diameters of the feeding vessel (DO = 0.84 cm) and stenotic vessel (do = 0.24 cm) are represented on each graph as dashed and dotted lines, respectively. Variations of measurement of the diameter of the feeding vessel (bold line) and the diameter of the stenotic vessel (lighter line) are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), and (c) filter for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); and transverse view, peak systolic velocity of 1 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|).

 


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  		Figure 6c. Graphs show results obtained with the high-frequency transducer. Baseline instrument settings are as follows: gain, 30%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%; and insonation angle, 60°. The reference diameters of the feeding vessel (DO = 0.84 cm) and stenotic vessel (do = 0.24 cm) are represented on each graph as dashed and dotted lines, respectively. Variations of measurement of the diameter of the feeding vessel (bold line) and the diameter of the stenotic vessel (lighter line) are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), and (c) filter for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); and transverse view, peak systolic velocity of 1 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|).

 


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  		Figure 6d. Graphs show results obtained with the high-frequency transducer. Baseline instrument settings are as follows: gain, 30%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%; and insonation angle, 60°. The reference diameters of the feeding vessel (DO = 0.84 cm) and stenotic vessel (do = 0.24 cm) are represented on each graph as dashed and dotted lines, respectively. Variations of measurement of the diameter of the feeding vessel (bold line) and the diameter of the stenotic vessel (lighter line) are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), and (c) filter for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); and transverse view, peak systolic velocity of 1 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|).

 


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  		Figure 7a. Graphs show results obtained with the low-frequency transducer. Baseline instrument settings are as follows: gain, 60%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%; and insonation angle, 60°. The reference diameters of the feeding vessel (DO = 0.84 cm) and stenotic vessel (do = 0.24 cm) are represented on each graph as dashed and dotted lines, respectively. Variations of measurement of the diameter of the feeding vessel (bold line) and the diameter of the stenotic vessel (lighter line) are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), (c) and filter, for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); transverse view, peak systolic velocity of 2 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|). Note that measurement data of the diameter of the normal vessel from the transverse view have been excluded.

 


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  		Figure 7b. Graphs show results obtained with the low-frequency transducer. Baseline instrument settings are as follows: gain, 60%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%; and insonation angle, 60°. The reference diameters of the feeding vessel (DO = 0.84 cm) and stenotic vessel (do = 0.24 cm) are represented on each graph as dashed and dotted lines, respectively. Variations of measurement of the diameter of the feeding vessel (bold line) and the diameter of the stenotic vessel (lighter line) are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), (c) and filter, for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); transverse view, peak systolic velocity of 2 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|). Note that measurement data of the diameter of the normal vessel from the transverse view have been excluded.

 


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  		Figure 7c. Graphs show results obtained with the low-frequency transducer. Baseline instrument settings are as follows: gain, 60%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%; and insonation angle, 60°. The reference diameters of the feeding vessel (DO = 0.84 cm) and stenotic vessel (do = 0.24 cm) are represented on each graph as dashed and dotted lines, respectively. Variations of measurement of the diameter of the feeding vessel (bold line) and the diameter of the stenotic vessel (lighter line) are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), (c) and filter, for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); transverse view, peak systolic velocity of 2 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|). Note that measurement data of the diameter of the normal vessel from the transverse view have been excluded.

 


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  		Figure 7d. Graphs show results obtained with the low-frequency transducer. Baseline instrument settings are as follows: gain, 60%; pulse repetition frequency, 1,500 Hz; filter, low; color versus echo priority, 100%; and insonation angle, 60°. The reference diameters of the feeding vessel (DO = 0.84 cm) and stenotic vessel (do = 0.24 cm) are represented on each graph as dashed and dotted lines, respectively. Variations of measurement of the diameter of the feeding vessel (bold line) and the diameter of the stenotic vessel (lighter line) are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), (c) and filter, for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); transverse view, peak systolic velocity of 2 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|). Note that measurement data of the diameter of the normal vessel from the transverse view have been excluded.

 


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  		Figure 8a. Graphs show evaluation of the degree of stenosis. The reference degree of stenosis (SO; 71%) is represented as a dashed and dotted line. Variations of the estimated degree of stenosis 3.0 cm deep (lighter line) and 11.5 cm deep (bold line), calculated for each situation reported in Figures 6 and 7, are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), (c) and filter, for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); transverse view, peak systolic velocity of 2 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|). Note that the estimation of degree of stenosis from the transverse view with the low-frequency transducer is not represented because measurement data of the diameter of the normal vessel have been excluded.

 


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  		Figure 8b. Graphs show evaluation of the degree of stenosis. The reference degree of stenosis (SO; 71%) is represented as a dashed and dotted line. Variations of the estimated degree of stenosis 3.0 cm deep (lighter line) and 11.5 cm deep (bold line), calculated for each situation reported in Figures 6 and 7, are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), (c) and filter, for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); transverse view, peak systolic velocity of 2 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|). Note that the estimation of degree of stenosis from the transverse view with the low-frequency transducer is not represented because measurement data of the diameter of the normal vessel have been excluded.

 


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  		Figure 8c. Graphs show evaluation of the degree of stenosis. The reference degree of stenosis (SO; 71%) is represented as a dashed and dotted line. Variations of the estimated degree of stenosis 3.0 cm deep (lighter line) and 11.5 cm deep (bold line), calculated for each situation reported in Figures 6 and 7, are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), (c) and filter, for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); transverse view, peak systolic velocity of 2 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|). Note that the estimation of degree of stenosis from the transverse view with the low-frequency transducer is not represented because measurement data of the diameter of the normal vessel have been excluded.

 


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  		Figure 8d. Graphs show evaluation of the degree of stenosis. The reference degree of stenosis (SO; 71%) is represented as a dashed and dotted line. Variations of the estimated degree of stenosis 3.0 cm deep (lighter line) and 11.5 cm deep (bold line), calculated for each situation reported in Figures 6 and 7, are represented when one of the following settings was changed: (a) gain, (b) pulse repetition frequency (PRF), (c) and filter, for the longitudinal view, peak systolic velocity of 1 m/sec ({triangleup}); longitudinal view, peak systolic velocity of 2 m/sec ({circ}); transverse view, peak systolic velocity of 2 m/sec (x). (d) Gain setting was changed for the longitudinal view and peak systolic velocity of 2 m/sec for insonation angles of 40° ({lozenge}), 60° (*), and 90° (|). Note that the estimation of degree of stenosis from the transverse view with the low-frequency transducer is not represented because measurement data of the diameter of the normal vessel have been excluded.

 
When the high-frequency transducer was used, no significant intra- or interobserver variability was observed (P = .9 and P = .57 for the measurement of the apparent lumen of normal and stenotic vessels, respectively). When the low-frequency transducer was used, a significant intraobserver variability was noted in one series (transverse view, insonation angle of 60°, peak systolic velocity of 2 m/sec; P = .03 and P = .08 for the measurement of the apparent lumen of normal and stenotic vessels, respectively). This series was then excluded from the analysis.

Results with the High-Frequency Transducer
The results of the measurements of the apparent lumen of normal and stenotic vessels are summarized in Figure 6, and the statistical analysis is presented in Table 1.


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  		TABLE 1. Results of Statistical Analysis with the High-Frequency Transducer
 
Measurements of the apparent lumen of stenotic vessels were generally slightly overestimated. Gain was the most significant factor, whereas velocity affected the results at high levels of gain—60% and 70%—where the measurements were partially affected by the encoding of vibrations around the vessels. Angle, view, pulse repetition frequency, and filter did not significantly affect the results.

Measurements of the apparent lumen of normal vessels on power Doppler US images were likely to be underestimated, except at high gain levels. Significant factors included gain, pulse repetition frequency, filters, and view.

As a result, the degree of stenosis was always underestimated, most values progressively decreasing from 65% and 48%, as compared with a known value of 71% (Fig 8). Evaluating the degree of stenosis was significantly affected by pulse repetition frequency, gain, velocity, and filter (Table 1).

Results with the Low-Frequency Transducer
The results are summarized in Figure 7, and the statistical analysis is presented in Table 2.


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  		TABLE 2. Results of Statistical Analysis with the Low-Frequency Transducer
 
Measurements of the apparent lumen of stenotic vessels were markedly overestimated. Gain, angle, and velocity were significant factors. Measurements of the apparent lumen of normal vessels were strongly affected by the changes of all the tested factors—angle, gain, velocity, pulse repetition frequency, and filters.

Some common findings were observed for the apparent lumen of both normal and stenotic vessels. (a) Both prominently increased when gain increased, with values at least twice normal for gain levels of 70%–80%. (b) For a peak systolic velocity of 1 m/sec, no signal was detected for gain levels below 50%, but for a peak systolic velocity of 2 m/sec, signal was detected for gain levels higher than 30%. (c) The transverse view appeared to be slightly more sensitive because some signal was detected with a gain level of 20%. (d) Pulse repetition frequency and filtering had a marked effect only on measurement of the apparent lumen of normal vessels, which dramatically decreased for high values of pulse repetition frequency and for increasing filtering. (e) The higher the insonation angle, the lower the value for the apparent lumen of normal and stenotic vessels.

The degree of stenosis was highly underestimated and showed considerable variations; in some cases, it was negative because the stenotic segment appeared larger than the feeding vessel (Fig 8). It was significantly affected by velocity, filter, and pulse repetition frequency (Table 2).

Comparison of Results Obtained with the Two Transducers for the Evaluation of the Stenosis
The degree of stenosis was significantly underestimated with the high- and the low-frequency transducers (P < .001 for both transducers). In addition, there was a difference between the two probes, the high-frequency transducer providing significantly better results than the low-frequency one (P < .001).


   DISCUSSION
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
REFERENCES
 
For the present study, we constructed a stenosis phantom without walls, which is assumed to reduce the highly attenuating wall and the impedance mismatches between the vessel lumen and tissue (19). We used a computer-monitored pump system to generate flows as similar as possible to those found in the internal carotid and renal arteries; velocities of 1 m/sec were considered normal, whereas velocities of 2 m/sec are usually detected in significant internal carotid and renal arterial stenoses (14,20).

Using a cardiac analysis system and analyzing jet flow areas, Jain et al (16) previously showed that the detection of low velocities with power Doppler US depends on the combination of instrument settings used, including pulse repetition frequency and filter. Our study results confirm the major role of technical settings in the imaging quality of the power mode.

During exploration of an internal carotid arterial stenosis, the level of backscattered signal is usually high because the vessel is superficial and a high frequency is used. With the experimental conditions of the present study, which only mimic and do not duplicate in vivo conditions, optimal settings were a gain of 30%, pulse repetition frequency below 1,500 Hz, and low filter level. With this range of settings and a longitudinal view, the velocity and the insonation angle did not contribute significantly to the measurements. However, with these optimal conditions, the degree of stenosis was underestimated by a factor of approximately 10%. These findings are consistent with the results obtained by Guo et al (15) who, using three-dimensional power Doppler US software, a linear 5-MHz probe, and a phantom, demonstrated that power Doppler US can be used to quantify arterial stenosis with an overall precision of 7%. However, these authors did not mention the depth of the stenosis they scanned.

The level of signal backscattered from the renal artery is much lower than that from the internal carotid artery in most patients; because the vessel is deeper, the signal is more attenuated, and a low frequency is used. When the low-frequency transducer was used in our model, power Doppler US results depended on most tested factors. It was difficult to define an optimal combination of settings: using a gain of 50%–60%, pulse repetition frequency of 1,500 Hz, and medium filtering appeared to give the best compromise, but the degree of stenosis was highly underestimated. However, a significant improvement of sensitivity to flow was observed for higher velocities, which could be useful to better evaluate stenotic arteries, in which blood flow is accelerated. As previously reported in phantom studies (4,15), a slight dependency of power Doppler US on the angle of incidence was confirmed in our study. This is likely related to wall filtering, even set at the lowest value (4,15).

We observed a better correlation between the experimental degree of stenosis and the estimated one from power Doppler US with the high-frequency probe (P < .001). This can be explained by the better transverse resolution of this transducer (color Doppler US frequency of 6.0 MHz) compared with the other one (color Doppler US frequency of 2.8 MHz). The best transverse resolution being 0.3 mm, the diameter measured at power Doppler US is likely to be an overestimation of the lumen of the stenotic vessel.

Another factor that explains the underestimation of the feeding vessel is the dependency of the power Doppler US signal on flow velocity distribution. Flow profile will gradually decrease from the center toward the edge of the vessel. This dependency is evidenced in Figure 4, which plots the estimated diameter of the feeding vessel as a function of the system pulse repetition frequency. As the pulse repetition frequency increases, the measured diameter decreases, and the system becomes less and less sensitive to the flow velocity. This results from the response of the wall filter used in the power Doppler US system. Increasing wall filtering to minimize motion artifacts or perivascular vibration encoding also contributes to increasing the risk for rejecting low-flow signal and making the power Doppler US evaluation of these vessels less accurate.

Gradation of the stenosis has low reliability, since it is subject to errors made in the measurement of the apparent lumen of both the stenotic and normal vessels. Power Doppler US cannot be used to measure stenoses accurately and mostly consistently leads to underestimation.

There are some limitations to the present study. Using a color versus echo priority of 100%, we observed artifacts due to encoding of motion of hyperechoic structures such as borders of feeding and stenotic vessels, especially when the gain level was high. The reduction of the color priority is likely to improve the accuracy of the method. However, the appropriate threshold is difficult to estimate and standardize, because the echogenicity of the phantom or arterial wall depends on many factors, including frequency, insonation angle, and degree of wall calcification.

In each series, we changed one setting, keeping the others constant, to identify the effect of each factor. Further studies may be useful to know whether combining two settings having opposite effects may improve the accuracy of power Doppler US; for instance, increasing gain and filtering may allow more accurate results in some conditions.

For the experiments with the low-frequency probe, the US gel pad with the stenosis was positioned at 11.5 cm deep from the transducer by inserting a water path between the probe and the flow tube. This condition is obviously different from an in vivo situation such as that in the renal artery because of the marked difference of attenuation between tissues and water. However, low to medium gain yielded an image brightness similar to the one obtained in vivo. Given the linearity of the US system over a large range of transmit power and receive gain, the experimental setup we considered provides a reasonable framework for this particular study.

These limitations apply to the current state of US systems. As more progress continues to be made toward higher resolution and more sensitivity, power Doppler US system performance is expected to become more adequate for reliable assessment of the degree of vessel stenosis.

Finally, our phantom model does not completely simulate in vivo conditions but only mimics them. It did not allow the study of some problems encountered when evaluating arterial stenosis in patients, including motion artifacts, tortuous or eccentric morphology of feeding vessels and stenoses, shadowing by calcified atheroma, and attenuation and scattering due to overlying tissues.

Practical application: Physicians should be aware of the advantages and limitations of power Doppler US when evaluating the morphology of normal and stenotic arteries, especially when measuring diameters or performing three-dimensional reconstruction (10). The degree of stenosis is more or less underestimated in all cases. The lower the frequency, the higher the risk of error, because an increasing number of factors plays a significant role.


   ACKNOWLEDGMENTS
 
The authors are grateful to George A. Taylor, MD, and Harriet J. Paltiel, MD, of the Children’s Hospital, Boston, Massachusetts, for their help in reviewing the manuscript.


   FOOTNOTES
 
9*. Vascular system, location unspecified Back

Author contributions: Guarantor of integrity of entire study, M.C.; study concepts, M.C., P.P.; study design, M.C., D.W., S.B.; definition of intellectual content, M.C., P.P.; literature research, M.C.; experimental studies, M.C., D.W.; data acquisition, M.C., D.W.; data analysis, M.C., D.W.; statistical analysis, S.B.; manuscript preparation, M.C., D.W., S.B.; manuscript editing, M.C.; manuscript review and final version approval, all authors.


   REFERENCES
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 

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