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Doppler Echocardiography: Principles and Applications
The Doppler Principle Using the principle first delineated by the physicist Johann Christian Doppler,53 one can use ultrasound to determine the velocity and direction of blood flow by measuring the change in frequency produced when sound waves are reflected from red blood cells.54,55 In this way, information regarding the presence, direction, velocity, and turbulence of blood flow can be acquired by cardiac ultrasound. The Doppler principle states that when a sound (or light) signal strikes a moving object, the frequency of that signal will be altered, and the increase or decrease in frequency will be proportional to the velocity and direction at which the object is moving. This is illustrated in Fig. 15–24. If a stationary transducer at the apex emits a sound wave with a transmitted frequency of fo and the wave is reflected by nonmoving red blood cells (RBCs) in an isovolumic phase of the cardiac cycle, then the received frequency fr will be identical to fo. If the signal is reflected by RBCs that are moving toward the transducer, as through the MV in diastole, the returning waves will be compressed so that fr will be greater than fo. Conversely, if the target RBCs are moving away from the transducer, as in the outflow tract in systole, the returning sound waves will be elongated and the received frequency will be decreased. Of importance, the magnitude of change in the received frequency is directly related to the velocity at which blood is flowing toward or away from the transducer.55 If the velocity of sound and the angle
where fd is the frequency shift recorded, fo the transmitted frequency, and c the velocity of sound. Note that the denominator is doubled because the sound wave does not originate with the RBC but must travel back and forth from the transducer. By measuring Doppler shift frequencies, the velocity and direction of blood flow can be calculated, displayed, and recorded. The angle between the direction of blood flow and the course of the sound beam is a most important factor in Doppler ultrasound (Fig. 15–25). Velocity is a vectorial entity, having magnitude and direction, and Doppler will detect only those velocities parallel or near parallel to the interrogating signal. Since the relationship between velocity and the angle is a cosine function and the cosine of angles up to 20 degrees is 0.9, little error is introduced within this range.55 Because the processor that calculates blood velocity assumes that the angle is 0 degrees, however, considerable errors occur when it is greater than 20 degrees. Moreover, the angle of incidence in 3D space usually cannot be determined with certainty from 2D echocardiographic images. Therefore, in order to obtain accurate velocity determination by Doppler, it is crucial to position and direct the transducer so that the beam is as parallel to flow as possible. In clinical use, the frequency of transmitted ultrasound is in the range of 2 to 7 MHz, the velocity of sound in tissue is approximately 1540 m/s, and the Doppler shift frequency is relatively small (approximately 1 to 4 kHz) as compared with the transmitted frequency. Because the Doppler shift frequencies are in the audible range, a speaker integrated into the Doppler echocardiography system can present them as an audible signal. Normal signals are tonal or musical. The Doppler shift can also be presented graphically to provide a hard copy printout and enable measurement.
Figure 15–26 shows the typical graphic pulsed Doppler pattern of normal systolic blood flow through the RV outflow tract into the PA, with flow velocity on the y axis and time on the x axis. The location and size of the area from which Doppler recordings are derived is determined by the operator by positioning a sample volume on the echo image. The absence of flow is represented by the zero or no-flow line, termed the baseline. By convention, flow toward the transducer is displayed above the baseline and flow away from the transducer is displayed below the baseline. The velocities above and below baseline represent flow toward or away from the transducer, not forward or backward in the circulation. Because of the effects of viscous friction, the sample volume almost invariably includes RBCs flowing at slightly different velocities. Even normal laminar blood flow in the great vessels varies in velocity across the lumen, as RBCs in the center of the vessel move at higher velocity than those exposed to viscous friction at the wall; this creates a parabolic rather than a flat flow profile. Therefore, any returning Doppler shifted signal contains a spectrum of velocities, each of which can be displayed by means of fast Fourier transform analysis. The graphic output of the Doppler signal displays the range of velocities within the sample volume site at any time in gray scale and the number of RBCs moving at any velocity as relative intensity. Normal laminar flow is characterized by a uniformity of velocity and direction of individual RBCs, and therefore a narrowly dispersed signal, while disturbed or turbulent flow is manifest by marked variability in velocity and direction and therefore a broad signal, which is multitoned, dissonant, and harsh.
Echographs have now been modified to enable recording of the low-velocity, high-amplitude Doppler signals produced by moving tissue as well as those of RBCs. The ability to assess tissue velocity provides an evaluation of transmural rate of contraction and relaxation.56 Also, Doppler tissue recordings permit assessment of regional function and appear to be quite useful in the assessment of diastolic function (see below).57,58 Finally, Doppler tissue recordings provide the basis for the derivation of regional strain measurements by echocardiography. Such measurements are independent of overall cardiac motion and passive motion due to tethering to adjacent myocardium and therefore enable the most accurate assessment of myocardial contractile performance. Continuous- and Pulsed-Wave Doppler Time-velocity spectral recordings of blood flow are generally obtained with two types of Doppler interrogation: continuous- and pulsed-wave (Fig. 15–27).59,60 In the continuous-wave (CW) mode, sound waves are both transmitted and received continuously. This requires two piezoelectric crystals in each transducer, one for transmitting and one for receiving. Because all flow velocities along the beam are recorded, CW Doppler cannot define individual signals at specific distances from the transducer—a problem referred to as range ambiguity. CW Doppler, however, has no upper limit of velocity that can be accurately recorded. Thus, a CW Doppler beam can accurately measure the direction and velocity of overall flow but cannot discern the precise site of origin of individual components within the signal (Fig. 15–28B).
The problem of range ambiguity can be overcome by pulsed-wave Doppler. In this mode, short bursts of signal are transmitted from the transducer at a given pulse-repetition frequency (PRF). The instrument then receives the signal for only a brief period—an interval that corresponds to the time required for sound energy to travel and return from a specific site along the beam path. In practice, the operator selects the location at which flow is to be examined by positioning a sample volume, and the instrument determines the period during which to receive the incoming reflected frequencies. With pulsed-wave Doppler, only a single piezoelectric crystal is needed and flow can be recorded in one small area within the heart or vasculature.59,60 Unfortunately, pulsed Doppler techniques employ intermittent sampling and are therefore susceptible to a problem of range ambiguity referred to as aliasing.61 By definition, aliasing is the erroneous representation of flow in the direction opposite to that in which it is actually occurring. To correctly record the velocity of blood flow by pulsed Doppler, the PRF must be at least double the Doppler shift frequency, a value known as the Nyquist limit. If the blood flow examined is of very high velocity or far from the transducer (requiring a long transit time), it may necessitate an unobtainably high PRF. In such cases, aliasing will occur as Doppler signals that depict flow at high velocity in ambiguous or opposite directions compared to actual flow (Fig. 15–28A). An intermediate mode between pulsed and CW methods, high-PRF Doppler, is also available. This mode enables higher-velocity recordings to be obtained at a compromise of depicting two to four sample sites simultaneously. Color-Flow Doppler The major limitation of pulsed and CW Doppler (sometimes referred to as spectral Doppler) is that no spatial information regarding the size, shape, and 2D direction of flow is provided. An extension of pulsed-wave Doppler techniques, color-flow Doppler (CFD), provides real-time M-mode or 2D imaging of blood flow by presenting the velocity and direction of RBC movement as shades of color superimposed upon gray-level 2D tissue structure. Standard pulsed Doppler yields flow signals from a single site along a single scan line. In CFD, rapid pulsed-wave interrogations are performed at multiple sites for multiple scan lines to create a spatially correct and dynamic display of moving blood within the heart and vasculature (Fig. 15–29). Doppler signals are presented as colors assigned to individual sites (Fig. 15–30). Blood flow moving toward the transducer is displayed in red, flow away from the transducer is displayed in blue, and increasing velocity is depicted in brighter shades of each color. The variance within each signal is calculated as a statistical marker of turbulence and is presented by adding green to the image (Fig. 15–31). Therefore, turbulent flow jets appear as a mosaic mix of colors. CFD also can be superimposed onto M-mode tracings (Fig. 15–32), often termed M/Q imaging, and is helpful in clarifying the timing of flow phenomena. Given the time constraints imposed by collecting the large volume of data required by CFD, velocity estimates are performed by autocorrelation techniques that are less accurate than fast Fourier transform analysis. Nevertheless, CFD technology is a major advance that has improved the rapid detection of cardiac pathology, especially valvular regurgitation and intracardiac shunts.
Normal and Abnormal Flow Dynamics The clinical application of Doppler recordings is based on the fundamental differences between normal and disturbed blood flow. Normal flow is laminar, with all RBCs exhibiting the same velocity and direction of flow. Although some abnormalities, such as atrial septal defects, involve laminar flow, most pathologic conditions involve disturbed or turbulent flow and share a common hydrodynamic basis for the resultant flow dynamics. Specifically, nearly all circulatory disturbances (stenosis, regurgitation, shunt) involve blood flow from a high-pressure chamber to a lower-pressure chamber through a restricted orifice.55 Aortic valve disease is a perfect example. Aortic stenosis is a forward flow disturbance in which turbulent blood travels from a high-pressure LV to a lower-pressure aorta through a restricted aortic orifice in systole. Aortic regurgitation (AR) is a retrograde flow disturbance in which turbulent blood regurgitates from a high-pressure aorta to a lower-pressure left ventricle through a small regurgitant orifice in diastole. In each case, the pressure gradient results in a high-velocity jet coursing through a restricted orifice, reaching its maximal velocity at a site just distal to the orifice, designated the vena contracta, at which time shear forces produce vortices resulting in flow of varying direction and velocity (Fig. 15–33). In each case, the velocity of the jet is related to the pressure gradient across the orifice. Thus, the hallmark of disturbed flow is a very high velocity jet with adjacent vortices of varying direction and velocity of flow. On pulsed Doppler recordings, these hemodynamic abnormalities cause broadening of the spectral signal and aliasing. On CW recordings, high velocity represents the primary abnormality. By color-flow imaging, the disturbance is manifest by the increased variance and higher velocities in the signal. With any of these techniques, of course, inappropriate timing of flow serves to highlight the abnormality (e.g., high-velocity LA flow during systole in mitral regurgitation).
The Standard Doppler Examination A clinical Doppler examination must be performed with full consideration of the three different Doppler modalities available, the types of information each can provide, the multiple sites for flow interrogation, and the spectrum of pathologic lesions that produces flow disturbances. In light of these considerations, it is understandable that the Doppler examination may not be as standardized as the format for 2D cardiac imaging. However, the vast majority of echocardiographic examinations include screening for flow disturbances by CFD. Since Doppler signals are best recorded with the ultrasound beam parallel to flow, screening is typically performed in long-axis or apical views. Any flow disturbances visualized are subsequently examined by CW spectral recordings and, in most laboratories, by pulsed-wave Doppler. Although CW examination is typically reserved for flow disturbances, pulsed-wave Doppler may also be of value in quantifying flow dynamics in the setting of laminar flow. In this regard, pulsed Doppler recordings obtained at the mitral, tricuspid, and aortic valvular orifices, PA, and pulmonary veins constitute part of a standard echocardiogram in many laboratories (Figs. 15–26 and Figs. 15–34, 15–35, 15–36, and 15–37).
The normal Doppler examination is characterized by uniformity of flow velocity and the absence of high-velocity turbulent flow. CFD recordings demonstrate laminar flow through the atrioventricular valves in diastole and the semilunar valves in systole. Since the Doppler examination is usually performed with a long-axis or apical transducer orientation, diastolic filling is characteristically encoded in red and ejection in blue (Fig. 15–30). Color aliasing is often observed at the levels of the mitral annulus and LV outflow tract as an abrupt change from bright red to bright blue or vice versa, usually in the center of the flow stream. Pulsed Doppler recordings of transmitral flow velocities are often recorded at the level of both the leaflet tips and annulus. Velocities are higher at the tips, while recordings at the annulus offer the ability to calculate flow through a cross-sectional area that is relatively uniform throughout the cardiac cycle. A sample volume positioned in the right upper pulmonary vein reveals systolic (S) and diastolic (D) flow velocities of nearly equal magnitude followed by a short, low-velocity reversal of flow into the pulmonary veins following atrial contraction (A) (Fig. 15–36). Flow in the LV outflow tract and aortic annulus area is characterized by a progressive increase of velocity peaking in early systole, followed by a more gradual deceleration of flow (Fig. 15–35). Minimal if any flow velocities are detected in the MV orifice and LV outflow tract in systole and diastole, respectively, in normal examinations. Examinations of the tricuspid and pulmonary valves give qualitatively similar results to those of the mitral and aortic valves (Figs. 15–26 and 15–37). Normal values for forward flow velocity are given in Table 15–4. As can be seen, velocity in normal individuals is highest in the aorta and is less than 2 m/s.62 Other commonly made measurements include the acceleration time (from the beginning of flow to peak velocity of flow in the ascending aorta or PA); and the deceleration time, from LV inflow peak E-wave velocity extrapolated to baseline zero velocity. Doppler Assessment of Diastolic Function In the last decade, there has been a great deal of interest in using mitral inflow velocity patterns to evaluate LV diastolic properties.63–69 Transmitral filling velocities reflect the pressure gradient between the LA and LV during diastole64 (Fig. 15–34). In early diastole, pressure in the LV normally falls below that in the LA, producing an increase in velocity due to rapid transmitral inflow (E wave). Flow decelerates as the pressures equilibrate in middiastole. In late diastole, LA contraction restores a small gradient, causing transmitral flow to accelerate to a second peak (A wave) that is of less magnitude than the E wave. In individuals in whom early LV relaxation is impaired, the transmitral pressure gradient is blunted, resulting in a decrease in both the velocity of early filling and rate of E-wave deceleration65,67 (Fig. 15–38A). Conversely, in patients with marked increases of LA pressure and LV stiffness, early diastolic filling velocities are high, deceleration is rapid, and late filling following atrial contraction is markedly reduced. This is the so-called restrictive pattern of LV filling (Fig. 15–38B). Accordingly, an E-wave velocity that is substantially less than the A-wave velocity and is accompanied by a prolonged deceleration time represents evidence of impaired early diastolic relaxation by Doppler, while an increased E-wave velocity and decreased A-wave velocity (E/A ratio greater than 2.5 or 3 to 1) accompanied by a diminished deceleration time (less than 160 ms) is indicative of a noncompliant LV with markedly elevated left atrial pressures.66,67 Although a restrictive pattern can be seen with restrictive cardiomyopathy or advanced LV dysfunction of any cause, it also occurs in pericardial disease.70 Of significance, a restrictive pattern of LV filling has been associated with an increased mortality rate in patients with advanced congestive heart failure,71 and persistence of this pattern despite changes in loading condition is an additional poor prognostic sign.72
These abnormal mitral inflow patterns can be clinically useful and, when they are markedly distorted, are generally reliable in identifying and characterizing diastolic dysfunction. Several variables other than diastolic function, however, are capable of influencing transmitral filling velocities. It has been shown that transmitral Doppler filling dynamics are affected by the age of the patient,73 changes in heart rate,74 respiration,75 and even the position of the Doppler sample volume within the MV orifice.76 Of greatest significance, transmitral inflow is very sensitive to loading conditions, and reductions in LV preload induced by nitroglycerin and/or lower-body negative pressure can induce a striking decrease in early transmitral filling velocities independent of changes in diastolic properties.77 The influence of LV loading upon transmitral filling is most striking when an increase in LA pressure due to cardiac dysfunction restores early diastolic filling velocities and obscures impaired relaxation, thus inducing "pseudonormalization." Therefore, because Doppler transmitral filling dynamics have many limitations in assessing diastolic function, particular filling patterns should not be interpreted as "pathognomonic" findings of diastolic dysfunction but rather as components of a complete clinical and echocardiographic evaluation. The recent addition of pulsed-wave tissue Doppler imaging (TDI) into the clinical arena has significantly enhanced the noninvasive assessment of diastolic function, especially when used together with transmitral and pulmonary venous PW Doppler. TDI measurements are obtained from the apical transducer position (four-chamber plane), with the sample volume placed on either the lateral or septal portion of the mitral annulus. Although TDI can assess systolic performance (i.e., by measuring systolic velocity of the mitral annulus toward the apex during systole), it is most often used to measure the motion of the annulus away from the transducer during diastole. The diastolic pattern is similar to the PW transmitral flow pattern, but the velocities are considerably less and in the opposite direction. The normal velocity of the early TDI motion (Em) is 12 cm/s or greater at the lateral annulus and 8 cm/s or more at the septal annulus. In addition, the ratio of transmitral E velocity to Em is normally in the range of 8 to 15. TDI can be combined with other Doppler modalities to assess diastolic function and estimate LV filling pressure. In a young, healthy individual, the E/A ratio is generally 1.5 to 2:1 (Fig. 15–34). Because of high LV compliance, the D velocity is greater than the S velocity in the pulmonary venous Doppler tracing (Fig. 15–36). With age, the LV compliance drops somewhat, so that by age 40 to 50, the S and D velocities are similar. As mentioned, TDI velocity is 12 cm/s or greater (Fig. 15–38C). In the setting of mild diastolic dysfunction, the E/A ratio is <1, the E deceleration time is prolonged ("relaxation abnormality,"Fig. 15–38A), and the pulmonary venous S wave is considerably larger than the D wave (S/D ratio >1) (Fig. 15–38D). Tissue Doppler imaging shows blunting of the Em wave with relative preservation of the later atrial component (Am) (Fig. 15–38E). As diastolic function worsens and LV filling pressures increase, a pseudonormal pattern occurs in the transmitral flow tracing (Fig. 15–39). Pulmonary venous and tissue Doppler imaging are especially helpful in this case: if the S/D ratio is <1 (except in the setting of a young individual), high LV filling pressure is likely present. Similarly, the presence of a low, blunted Em velocity in the setting of a "normal" transmitral E/A ratio strongly suggests diastolic dysfunction and elevated LV filling pressure. Of note, the Valsalva maneuver can be helpful in cases of pseudonormal transmitral flow patterns. In normal individuals, both E and A velocities drop to a similar degree with Valsalva. In pseudonormal cases, however, the drop in preload caused by the Valsalva maneuver changes the transmitral pattern to that of mild diastolic dysfunction (and the pseudonormal tracing changes to that of a "relaxation" abnormality).
In severe diastolic dysfunction, transmitral flow demonstrates a "restrictive" pattern, with an abnormally high E/A ratio and a markedly shortened E wave deceleration time (Fig. 15–38B). Concomitant pulmonary venous tracings show a very low S velocity and elevated D velocity (S << D) (Fig. 15–38F). In some cases, the pulmonary venous atrial reversal wave can be very prominent and prolonged (this is not a universal finding, as atrial systolic dysfunction can sometimes occur along with severe LV diastolic failure). In this regard, an abnormally prolonged duration of reversed pulmonary venous flow during atrial contraction (i.e., a longer duration than that of the forward transmitral flow during the A wave) accurately predicts elevated LV filling pressures.64 TDI in severe LV diastolic dysfunction shows marked blunting of both Em and Am velocities (Fig. 15–38G). An exception to this occurs in constrictive pericarditis, where early diastolic mitral annular motion is often preserved (as the myocardium is not inherently abnormal). Thus, when transmitral and pulmonary venous Doppler suggest severe diastolic dysfunction, a normal TDI pattern suggests constrictive rather than restrictive physiology. Finally, color M-mode imaging has been used to assess the velocity of propagation of the transmitral filling stream into the left ventricle. This technique appears helpful in distinguishing constrictive pericarditis from restrictive cardiomyopathy. Some reports have suggested that the velocity of propagation is preload-independent,58 but this has been questioned. Doppler Assessment of Systolic Function and Cardiac Output Although measurements of LV volume and ejection fraction can be obtained by 2D echocardiography, Doppler interrogation provides a unique and complementary noninvasive assessment of systolic function. Thus, LV systolic dysfunction often results in decreased aortic velocity and acceleration time. As discussed below, in the presence of mitral regurgitation (MR), the acceleration of the MR jet can provide information regarding contractile function.78 One of the most important applications of Doppler is in the calculation of the stroke volume.79 The theory involved is relatively simple. The volume of flow through any orifice or tube can be calculated as the product of the cross-sectional area through which flow occurs and the velocity of that flow (Fig. 15–40). Measurements of anatomic cross-sectional area can be derived from echocardiographic images, while velocity can be determined by Doppler. As the annulus of the aortic valve is nearly circular, its cross-sectional area can be estimated from a measurement of diameter as
Calculation of stroke volume by the Doppler method involves a number of assumptions. The orifice must be circular and constant in size, and the flow velocity must be uniform throughout the cross-sectional area. In addition, the angle between flow and the interrogating beam must be less than 20 degrees. Despite the uncertainty of these assumptions, Doppler-derived measurements of cardiac output and stroke volume have been shown to correspond well with thermodilution, Fick, and the angiographic calculations, though the correlation is not perfect.79–81 Theoretically, stroke volume can be calculated at any valve annulus.80,82 In clinical practice, however, this is not always possible (e.g., it is difficult to obtain an accurate diameter of the PA in every patient). Because the measurement of annular radius is squared in the computation of area, it is the most important source of error of Doppler stroke-volume analyses. Stroke-volume analysis through the mitral annulus is cumbersome; it is uncertain whether the mitral annulus is best described as a circle or an ellipse, and the cross-sectional area of the annulus probably changes slightly during diastole. Calculations using the tricuspid annulus are hampered by similar problems. Despite these limitations, measurements of stroke volume through the various cardiac valves are clinically useful and can be used to calculate pulmonary-to-systemic shunt ratios, regurgitant volumes,83,84 and orifice areas of stenotic valves by the continuity equation85,86 (see below). The Bernoulli Equation An important application of Doppler echocardiography is the calculation of pressure gradients within the cardiovascular system using a modification of the Bernoulli equation.87 This theorem states that the pressure drop across a discrete stenosis in the heart or vasculature occurs because of energy loss due to three processes: (1) acceleration of blood through the orifice (convective acceleration), (2) inertial forces (flow acceleration), and (3) resistance to flow at the interfaces between blood and the orifice (viscous friction).88 Therefore the pressure drop across any orifice can be calculated as the sum of these three variables (Fig. 15–41). In most clinical situations, the contribution of inertial forces and viscous friction are minimal and can be discounted. Since convective acceleration is determined by velocity, the pressure gradient can be calculated from the velocities of blood proximal to and at the level of an orifice as gradient = 4[(orifice velocity)2– (proximal velocity)2]. If the blood velocity proximal to the stenosis is low (<1.0 m/s), this term can be ignored as well. The resulting modified equation states that the pressure gradient across a discrete orifice is equal to four times the square of the peak velocity (V) through the stenosis (PG = 4V2).87,88
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between the direction of RBC flow and the beam path are known, then the velocity of the RBCs is described by the Doppler equation:
(diameter/2)2. The pulsed-wave Doppler envelope also can be recorded at the same level. The mean flow velocity through the orifice is calculated by integrating velocity over time (that is, by measuring the area under the Doppler curve). This velocity-time integral, often called the stroke distance, is then multiplied by the cross-sectional area at the level of the Doppler interrogation to obtain the stroke volume.
