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Christian Doppler, in his 1842 paper titled, “On the Coloured Light of the Binary Stars and Some Other Stars of the Heavens,” was the first to note that stars moving towards the earth emitted blue light while stars moving away from earth radiated red light. Thus, he postulated that the observed frequency of a wave depends on the relative speed of the source and the observer. Although Doppler himself never extended the principle to other natural phenomena, the common observation that the pitch of sound is different for a locomotive approaching the listener than one moving away led to a more widespread application of Doppler's initial observation.
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In clinical echocardiography, the Doppler technique depends upon an analysis of the frequency and wavelength of an emitted ultrasound beam. Frequency is defined as the number of waves passing though a certain point in 1 second, and is a fundamental property of the sound waves. It is expressed in units of hertz (Hz) and determines the resolution and the depth of penetration of the medium. However, the Doppler assessment of ultrasound waves depends upon not just the absolute emitted frequency, but the relative change in frequency as the sound waves are reflected back (by red blood cells) towards the transducer. The frequency of the reflected sound waves increases when the red blood cells are moving towards the transducer and decreases when red blood cells are moving away from the transducer (Figure 4–1). This relative change in the frequency, known as the Doppler shift, enables the echocardiographer to assess the direction, speed, and turbulence of blood flow, which in turn helps to objectively quantify intracardiac pressures and the severity of valvular stenosis and regurgitation.
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The Doppler shift is defined as the change in frequency of the reflected ultrasound waves and is described by the following mathematical relationship1:
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- Δf = 2ft × (v × Cosθ)/c
- Δf = Difference between the transmitted frequency (ft) and received frequency
- v = Velocity of red blood cells
- θ = Angle between the Doppler beam and the direction of blood flow
- c = Speed of ultrasound in blood (1540 m/s)
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When the Doppler beam is parallel to the direction of blood flow, the cosine of θ is 1 and the Doppler shift is most accurately calculated, but with increases in θ there is a progressive decrease in the measured Doppler shift. This underestimation of Doppler shift remains clinically insignificant until θ exceeds 20° and the associated percentage error exceeds 7% (Figure 4–2).2 When the Doppler beam is perpendicular to the direction of flow, the Doppler shift is recorded as zero since the cosine of 90 is zero. It is therefore essential to align the Doppler beam as parallel to the direction of blood flow as possible for an accurate measurement of speed and direction of flow.
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All ultrasound systems are equipped to create both a spectral display and an audio component of the beat-to-beat Doppler data during acquisition. The loudness and pitch of this audio signal varies with the strength and quality of the Doppler data, and experienced echocardiographers have been known to use the audio component alone to acquire and record the maximal velocity. The loudness/pitch of the audio signal can also be used to qualitatively diagnose the presence of stenosis; however, the information obtained from audio analysis is qualitative and subjective and should not be used as the sole means of quantification.1
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On the other hand, quantitative Doppler data on a spectral display provides a high degree of spatial and temporal resolution. Before the Doppler information can be displayed as distinct envelopes (velocity profiles over time) on the spectral display, the raw Doppler data undergo significant post-acquisition manipulation (demodulation, fast Fourier transformation, and chirp-Z transformation) to isolate the required frequencies.1 If the returning frequency is higher than the transmitted frequency, it is referred to as a “positive Doppler shift” (blood moving towards the transducer) and is displayed above the baseline. If the returning frequency is less than the transmitted frequency, it is called a “negative Doppler shift” (blood moving away from transducer) and is displayed below the baseline.
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It is important to remember that a simultaneous display of the two-dimensional (2D) image and Doppler information requires a time-share arrangement between two independent functions of the transducer. In the case of a continuous spectral display, the 2D image display is rapidly switched on and off, giving the impression of a continuous image alongside the spectral display. Thus, there is always some reduction in the quality of 2D and Doppler data when a simultaneous imaging mode is used.1
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Doppler can be used either in “pulse” or “continuous” mode to interrogate intracardiac flow patterns. Both these modalities have strengths and weaknesses that are suited for specific situations.
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Pulsed-Wave Doppler (PWD)
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PWD utilizes a single ultrasound crystal that alternates between transmission and reception. As a consequence, the maximal frequency shift (and thus maximum velocity) that can be measured is limited to one-half the pulse repetition frequency (also known as the Nyquist limit). When this maximal frequency shift is exceeded, aliasing or a wrapping around of the velocities is seen on the spectral Doppler display. In general, the maximum measurable velocity without aliasing with PWD is less than 2 m/s. The pulse repetition frequency (PRF) is directly related to the depth of placement of the sample volume, ie, the longer the time it takes for the sound to travel back to the transducer, the lower the PRF (Figure 4–3). An advantage of PWD is that the operator can select the specific site of measurement by manually positioning the sample volume during conventional 2D echocardiography, thus providing “range resolution” or the ability to identify the exact location of the recorded velocity (Figure 4–4). Simultaneous display of the location of the sample volume and the spectral Doppler data is also possible with some sacrifice of the quality of the 2D image as well as the Doppler signal. The distance from the transducer determines the size and three-dimensional shape of the sample volume. The farther the sample volume is placed from the transducer, the larger it becomes due to progressive divergence of the ultrasound beam from the transducer.
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During transmission of a PWD beam, the signals returning to the transducer are interpreted on the basis of the time it takes for them to return to the transducer, a process referred to as “time-gating.” During time-gating, signals returning after a specific time from a preselected depth are chosen for interpretation, and all other returning signals are selectively ignored by the ultrasound system. Time-gating is described by the following mathematical relationship:
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- d = Depth
- C = Speed of sound in blood
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High Pulse Repetition Frequency (HPRF) Doppler
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A technique to measure high velocities with PWD employs special transducers with the ability to emit multiple pulses. Since any emitted pulse does continue beyond the primary sampling depth (albeit weaker), the net effect is increased sampling at locations beyond the primary depth. Depending on the number of pulses, the Nyquist limit can be increased by a factor of 2 or more. The increase in sampling frequency is, however, at the expense of range ambiguity. The primary application of HPRF Doppler is in resolving the velocity of multiple high-velocity lesions that are in series (Figure 4–5).
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Continuous-Wave Doppler (CWD)
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In the continuous-wave mode, the ultrasound transducer utilizes two crystals—one to continuously send an ultrasound beam and another to constantly receive the reflected wave (see Figure 4–4). Since there is one crystal dedicated to receiving the reflection, the maximal frequency shift that can be detected is not limited by the pulse repetition frequency, and very high velocities can be recorded without aliasing. However, the CWD signal is not time-gated and measures the highest velocity in the path of the ultrasound beam, thus suffering from “range ambiguity” or the inability to discriminate the precise location of the highest recorded velocity. The velocity envelopes obtained with CWD are typically shaded, representing the multiple velocities measured in the course of the ultrasound beam.
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Color-flow Doppler is a pulsed ultrasound technique that color codes Doppler information and superimposes it on a 2D image, displaying information on the direction and the mean velocities of flow as color maps. Blood flow directed towards the transducer is commonly (but not exclusively) color-coded in shades of red, while blood flowing away from the transducer is color-coded in shades of blue. Lighter shades within each of these colors typically represent higher velocities. Color-flow Doppler uses packets of multiple pulses (3 to 20 per scan line), and therefore has a low temporal resolution. It has the characteristics of pulsed-wave Doppler (range discrimination and aliasing), but color-flow Doppler aliases at a lower velocity than traditional PWD because of the reduction in PRF associated with the simultaneous generation of a gray-scale and a color image.
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Doppler Tissue Imaging
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Conventional Doppler modalities are designed to calculate the intracardiac or intravascular blood flow velocities (ie, high-velocity and low-amplitude signals). Doppler tissue imaging (DTI) is geared to detect velocities of the actual myocardial tissue (ie, low-velocity, high-amplitude signals).3–5 Current echocardiography systems come equipped with presets for measuring DTI signals, and are generally set to measure myocardial velocities in the range of 0.2 to 0.4 cm/s and detect amplitudes greater than 20 dB. In contrast, the amplitude of the movement of red blood cells is generally less than 15 dB.6 The DTI signals can be displayed as a spectral tracing, in association with color-flow Doppler signals, or as color-encoded 2D signals.
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Two-dimensional and Doppler imaging can be used to obtain comprehensive hemodynamic data, but it should be remembered that these data, while providing a more objective measure of cardiac performance, should always be interpreted within the context of the patient's clinical condition.
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In circulation, the blood flow and velocity are phasic, ie, change throughout the cardiac cycle. A Doppler spectrum of the velocity of blood flowing through a conduit will yield a curve during systole that has velocity (cm/s) on the y-axis and time (s) on the x-axis. When this curve is integrated, it yields a velocity-time integral (VTI) in units of centimeters (velocity [cm/s] time [s] = VTI [cm]). The VTI is a manifestation of the stroke distance (ie, the distance the stroke volume travels over time during a single systolic ejection period). The product of VTI (cm) and cross-sectional area (CSA; cm2) will yield stroke volume (SV; cm3):
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Stroke Volume (SV) = Stroke Distance (VTI) × Cross-Sectional Area (CSA)
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Cardiac output can then be calculated as:
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Cardiac Output = Stroke Volume × Heart Rate
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Cardiac index can be calculated as:
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Cardiac Index = Cardiac Output/Body Surface Area
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The calculation of SV through a conduit assumes laminar flow, a flat flow profile in which velocities are uniform, constant diameter of the conduit, velocity measured at the same point as the diameter, velocity measured represents the mean velocity during the entire ejection period, and accurate measurement of the conduit diameter (a common source of error). Stroke volume can be measured at any conduit in the heart; however, the preferred site for such calculation is the left ventricular outflow tract (LVOT) (Figure 4–6). Other intracardiac sites such as the pulmonic and mitral valves have a dynamic conduit diameter during the cardiac cycle, and at times the velocity profile is parabolic, where velocities vary across the parabola.
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In echocardiography, it is assumed that a conduit or orifice, such as the LVOT, is circular. The CSA is calculated from the diameter measurements by using the equation for area of a circle:
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The diameter of the LVOT is calculated from the midesophageal long-axis window, and LVOT VTI is calculated from the deep transgastric window (Figure 4–7).
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Measurement of Intracardiac Shunts
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The ratio of pulmonary (Qp) to systemic (Qs) flow is a useful parameter to assess the magnitude of shunting, with a ratio greater than 1.5 generally considered significant. Qp/Qs calculation requires calculation of the stroke volume at the right ventricular outflow tract (RVOT) and the left ventricular outflow tract (LVOT). Thus, for atrial or ventricular shunts:
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Qp/Qs = (CSARVOT × TVIRVOT)/(CSALVOT × TVILVOT)
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For shunts occurring through a patent ductus arteriosus:
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Qp/Qs = (CSALVOT × TVILVOT)/(CSARVOT × TVIRVOT)
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Assuming a constant flow of fluid through a conduit at a certain velocity, if there is a stenosis in the conduit, the velocity of fluid will increase at the site of stenosis to conserve flow. This concept is known as the continuity of flow and sometimes as the conservation of flow:
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FlowLarger Conduit = FlowStenosis
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As described above, constant flow (cm3/s) in a conduit is the product of cross-sectional area (CSA) of the conduit (cm2) and the average velocity of the fluid (cm/s). Thus,
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CSALarger Conduit × VelocityLarger Conduit = CSAStenosis × VelocityStenosis
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When three variables are known, the fourth is easily determined with this equation, commonly known as the continuity equation.
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In aortic stenosis, flow across the aortic valve is equal to the flow across the LVOT (Figure 4–8) and in order to determine the stenotic area, the continuity equation can be reordered as:
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Mitral valve area can also be similarly calculated, but it must be remembered that the transmitral flow must be the same as left ventricular SV, a condition that is met only in the absence of ventricular shunts and mitral and aortic regurgitation.
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The principles of flow conservation can also be applied to assess regurgitant volumes and fractions. For this calculation, one valve is considered the “reference valve,” but it should be remembered that the assumption that the mitral valve orifice is circular may not always be valid. In the absence of aortic stenosis or regurgitation (see Chapters 7 and 9), SVLVOT can be substituted for SVAortic Valve
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Velocity Acceleration
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When molecules move within a large cavity toward a small orifice, as in stenotic and regurgitant lesions, the velocity increases over a large area and the velocity profile is hemispherical, with the cavity of the hemisphere facing the orifice. The velocity over the surface of the hemisphere is the same (isovelocity), and because it is proximal to the orifice, the surface area is known as proximal isovelocity surface area (PISA). The product of the (iso)velocity (cm/s) and the surface area of the hemispherical velocity profile (cm2) yields the flow (cm3).
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If the flow approaching the orifice is examined with color-flow Doppler, with the color scale set so that the accelerated velocity exceeds the Nyquist limit, aliasing will take place and a semicircular shell of a contrasting color will appear to cap the orifice (Figure 4–9). To obtain this shell, the baseline for the color scale should be shifted in the direction of the jet of interest (eg, towards the transesophageal echocardiography [TEE] transducer for mitral regurgitation). The semicircular shell is in fact a hemisphere in three dimensions, and its surface area can be calculated as that of a sphere:
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The velocity at the surface of the hemispherical velocity profile is the Nyquist limit (aliasing velocity) on the color-flow Doppler scale. Thus,
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By the principle of conservation of flow, the flow through an orifice is the same as the flow where the PISA is located:
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In the setting of stenosis, the valve area can be calculated by reordering the equation as:
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For a regurgitant lesion, the calculation of the effective regurgitant orifice area (EROA) employs the peak velocity of the regurgitant jet so that:
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The PISA radius should be measured at the same time as the peak velocity of the jet, and this can be more readily accomplished using color M-mode imaging. PISA has been validated for mitral valve assessment, but is not commonly applied to TEE assessment of the aortic valve. PISA is performed in diastole and on the left atrial side for assessment of mitral stenosis severity, and during systole and on the left ventricular side for calculation of the EROA (Figure 4–10). Summarizing, the assessment of the mitral valve:
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True hemispheric shells require a flat valve surface area, and because the mitral valve surface area is not flat, an angle correction term is sometimes used to increase the accuracy of volumetric flow assessment:
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where α is the angle subtended by the mitral leaflets (see Figure 4–9). However, this technique of angle correction is limited by the inability to readily measure α with existing ultrasound system software.
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Measurement of Pressure
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A Dutch-Swiss mathematician, Daniel Bernoulli, in 1798 proposed the principle that the total energy in a steady flowing fluid column remains constant, such that when the velocity (kinetic energy) increases, it is accompanied by a simultaneous decrease in pressure (potential energy). This principle helped to explain the development of a pressure gradient across a narrowing, ie, there is an increase in distal velocity with a simultaneous drop in pressure (Figure 4–13). The Bernoulli equation can be applied to compressible and noncompressible fluids or even gases, and is expressed as:
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- P1 – P2 = Pressure difference between the two locations
- ρ = Mass density of blood (gm/cm3)
- V1 = Velocity proximal to stenosis (m/s)
- V2 = Velocity at vena contracta (m/s)
- dv/dt = Acceleration
- s = Distance over which flow accelerates
- R = Viscous resistance
- μ = Viscosity
- v = Velocity of blood flow (m/s)
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During routine clinical echocardiography, acceleration at peak velocities (dv/dt) is zero and viscous acceleration is negligible, and hence both terms can be ignored. One-half of the mass density of blood (1/2ρ) is equal to 4 (after correction for units of measure), hence the Bernoulli equation can be modified as:
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The proximal velocity (V1) is also generally very low and can often (but not always) be ignored, thus further simplifying the equation:
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Limitations of the Bernoulli Equation
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Significant error can be introduced into the calculation in specific situations when the assumptions inherent in simplifying the equation are violated. For example, flow acceleration (inertial force) can become significant with some prosthetic valves, where a greater than normal force is required to open the valve. Similarly, the presence of viscous friction is negligible with laminar flow, but should be accounted for in lesions with tubular obstructions greater than 4 cm in length and orifices less than 0.1 cm2. The simplified Bernoulli equation also ignores the proximal velocity (V1), but in conditions such as high cardiac output states, sub-aortic obstruction, significant aortic regurgitation, and intracardiac shunts, the proximal velocity (V1) can be significant (>1.5 m/s), leading to an overestimation of the pressure gradient if the modified (rather than the simplified) Bernoulli equation is not applied. Finally, alterations in the blood viscosity, such as an increase in hematocrit to 60%, may lead to an underestimation of gradients, as 1/2 ρ may be higher from an increase in the mass density of blood.
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Applications of the Bernoulli Equation
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Peak and Mean Gradients The Bernoulli equation is most commonly used to determine the peak instantaneous gradients across stenotic valves. The peak instantaneous gradient is measured with continuous-wave Doppler, and most echocardiography machines can automatically calculate the peak gradient by simply positioning the cursor at the highest point of the velocity envelope (Figure 4–14).
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Peak Instantaneous Gradient = 4 × (VPeak)2
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Mean pressure gradients are calculated as the average of multiple successive peak instantaneous peak gradients measured over time during the particular ejection phase (see Figure 4–14).
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Intracardiac Pressure Measurements Estimation of intracardiac chambers is one of the most common forms of application of the modified Bernoulli equation. This method requires the presence of a regurgitant jet or the presence of a shunt jet. The estimation of intracardiac pressures is performed in the following steps (Figure 4–15)2:
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Utilization of continuous-wave Doppler to measure the peak velocity of the jet
Conversion of peak velocity into pressure gradient with the Bernoulli equation
Estimation of pressure in the origination chamber (POC)
Estimation of pressure in the receiving chamber (PRC)
Calculation of pressures (Figure 4–15):
POC = 4v2 + PRC
PRC = POC − 4v2
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Using this methodology, numerous intracardiac pressures can be measured depending upon the presence or absence of regurgitation jets:
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Right Heart Pressures
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Right ventricular systolic pressure (RVSP) using tricuspid regurgitant (TR) jet
RVSP = 4 (VPeak TR)2 + Right Atrial Pressure (CVP)
In the absence of pulmonic stenosis or right ventricular outflow tract obstruction, RVSP is equal to pulmonary artery systolic pressure.
Right ventricular systolic pressure in the presence of a ventricular septal defect (VSD)
RVSP = Left Ventricular Systolic Pressure − 4 (VPeak VSD)2
Pulmonary artery mean pressures (PAMP) using pulmonary regurgitant (PR) jet
PAMP = 4 (VPeak PR)2 + Right Atrial Pressure (CVP)
Pulmonary artery diastolic pressures (PADP) using pulmonary regurgitant (PR) jet
PADP = 4 (VEnd-Diastolic PR)2 + Right Atrial Pressure (CVP)
Pulmonary artery systolic pressure (PASP) in the presence of a patent ductus arteriosus (PDA)
PASP = Systolic Blood Pressure − 4 (VPeak PDA)2
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Left atrial pressure (LAP) from a mitral regurgitant (MR) jet
LAP = Left Ventricular Systolic Pressure − 4 (VPeak MR)2
In the absence of aortic stenosis or left ventricular outflow tract obstruction, systolic blood pressure can be substituted for left ventricular systolic pressure.
Left atrial pressure in the presence of a patent foramen ovale (PFO)
LAP = 4 (VPeak PFO)2 + Right Atrial Pressure (CVP)
Left ventricular end-diastolic pressure (LVEDP) from an aortic regurgitant (AR) jet
LVEDP = Diastolic Blood Pressure − 4 (VEnd-Diastolic AR)2
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Measurement of Resistance
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Systemic vascular resistance (SVR) and pulmonary vascular resistance (PVR) are typically calculated from invasive hemodynamic measurements; however, Doppler echocardiography can provide a noninvasive assessment of vascular resistance. Units for measuring vascular resistance are dyne·s·cm−5, pascal seconds per cubic meter (Pa·s/m3), or mm Hg/L/min, which is referred to as a Wood unit (WU). The WU value is multiplied by 8 to convert to Pa·s/m3 or by 80 to obtain the value in dys·n·cm−5. Normal SVR values range from 10 to 14 WU, while a normal PVR is 1 WU.
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The ratio of peak mitral regurgitant velocity to the time-velocity integral of the LVOT flow (VMR/ TVILVOT) measured by Doppler echocardiography has been shown by Abbas et al7 to correlate positively with SVR measurements (WU) obtained invasively (r = 0.84, 95% CI = 0.7 to 0.92). Furthermore, a calculated ratio greater than 0.27 identified patients with elevated SVR (>14 WU) with 70% sensitivity and 77% specificity, whereas a ratio less than 0.2 had a 92% sensitivity and 88% specificity to identify SVR less than 10 WU.
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Similarly, Doppler echocardiography has been shown to provide a clinically reliable noninvasive method to determine PVR8:
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where VTR = peak tricuspid regurgitant velocity and TVIRVOT = right ventricular outflow tract time-velocity integral. Furthermore, the ratio of VTR/TVIRVOT compared favorably to invasive PVR measurements (r = 0.93, 95% CI = 0.87 to 0.96), and a ratio greater than 0.175 had a sensitivity of 77% and a specificity of 81% to determine PVR greater than 2 WU.8 Scapellato et al9 have also described a method of estimating PVR using the preejection period (PEP), acceleration time (AcT), and total systolic time (TT) derived from pulmonary systolic flow:
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PVR = −0.156 + {1.154 × [(PEP/AcT)/TT)]}
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Although these methods have the advantage of being simple and easily applicable, they may not be as reliable in patients with pulmonary arterial hypertension (PAH). More recently, Haddad et al10 reported that the ratio of systolic pulmonary artery pressure/(HR × TVIRVOT) correlated very well with invasive measurements of PVR indexed to body surface area (PVRI; r = 0.86; 95% CI = 0.76 to 0.92). A cutoff value of 0.076 provided a sensitivity of 86% and specificity of 82% to determine PVRI greater than 15 WU/m2. A cutoff value of 0.057 increased sensitivity to 97% but decreased specificity to 65%. Similarly, Kouzu et al11 showed that the ratio of the peak tricuspid regurgitant pressure gradient over the time–velocity integral of right ventricular outflow (PGTR/TVIRVOT) provided a reliable estimation of PVR over a wide range of PAH values and from various causes, including intracardiac shunts. In addition, a PGTR/TVIRVOT ratio greater than 7.6 was suggestive of poor prognosis for patients with PAH without an intracardiac shunt.
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Measurement of Contractility
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A relatively load-independent index of left ventricular systolic performance is peak dP/dt, or the maximum rate of rise of left ventricular pressure during systole. The echocardiographic method for deriving this parameter is based on the continuous-wave Doppler recording of the mitral regurgitant spectrum, wherein the time for velocity to rise from 1 m/s to 3 m/s is measured, and the pressure change from 1 m/s to 3 m/s is calculated by the Bernoulli equation as 32 mm Hg (4 × 32 − 4 × 12). dP/dt is then calculated with the following equation:
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Left Ventricular dP/dt = 32 × 1000/dt in Milliseconds
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Normal values for this parameter are 1610 ± 290 mm Hg/s. Although relatively afterload dependent, dP/dt is preload dependent.
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Right ventricular dP/dt may be calculated by using the continuous-wave tricuspid regurgitant spectrum in a manner analogous to the approach used for left ventricular dP/dt. Because even hypertensive right ventricular pressures are typically lower than those of the left ventricle, the convention is to make the calculation based on the rise in velocity between 1 and 2 m/s. The pressure change from 1 m/s to 2 m/s is calculated by the Bernoulli equation as 12 mm Hg (4 × 22 − 4 × 12). dP/dt is then calculated as:
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Right Ventricular dP/dt = 12 × 1000/dt in Milliseconds
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A value greater than 1000 mm Hg/s is generally associated with normal right ventricular function.
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There has been a lack of standardization of values for cardiac chamber quantification with echocardiography, particularly TEE. Comparison of TEE measurements with transthoracic echocardiography (TTE) standards has also added to the confusion and led to the impression that the measurements made during echocardiography are somehow less accurate than with other imaging modalities such as magnetic resonance imaging (MRI). However, despite differences in imaging planes, in general, the same range of normal values are recommended for both TEE and TTE.12
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To make reliable measurements with echocardiography, it is essential that all conditions be optimized for image acquisition, display, and archiving. The echocardiographer should make an effort to minimize translational movements of the heart, make adjustments to maximize image resolution, avoid apical foreshortening of the left ventricle, and optimize endocardial definition. Furthermore, it is important to identify systole and diastole with simultaneous display of the electrocardiogram (ECG) and to make measurements at the appropriate point in the cardiac cycle. For patients with arrhythmias, it is critical that measurements be averaged over multiple cardiac cycles.12
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End-diastole is identified temporally along the ECG tracing as the onset of the QRS complex. However, end-diastole can also be defined as the frame after mitral valve closure or as the frame with the largest cardiac dimension. End-systole is defined as the frame preceding mitral valve opening or the frame with the smallest cardiac dimension.12
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Quantification of Left Ventricle (LV)
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A variety of echocardiographic techniques have been proposed to quantify LV dimensions and volumes (Table 4–1). Of the many potential LV measurements, septal wall thickness (SWT), inferolateral (posterior) wall thickness (PWT), and internal dimensions during systole (LVIDs) and diastole (LVIDd) are the most clinically relevant. PWT and SWT are best assessed in the transgastric mid–short axis view (Figure 4–16) in diastole. LV diameters are ideally measured from the midesophageal two-chamber and the transgastric two chamber views (Figure 4–17), but care should be taken to avoid apical foreshortening. LVID is measured at the minor axis of the LV (ie, at the tips of the mitral valve leaflets), and the range for normal systolic and diastolic measures at this level are 3.3 ± 0.5 cm and 4.7 ± 0.4 cm, respectively (Tables 4–2, 4–3).12 While 2D or M-mode can be used to make these measurements, temporal resolution is better with M-mode leading to more accurate measurements. When M-mode is applied, the distance between leading edge echoes is measured.
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The current method of choice for LV volume measurement is the biplane method of disks (modified Simpson rule). It is based on modeling the left ventricle as a series of stacked cylindrical disks capped by an elliptical disk apex. The volume for each cylindrical disk is quantified by using the equation:
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V = (π × D1/2 × D2/2) × H
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where D1 and D2 are orthogonal diameters of the cylinder, and H is the height of the cylinder. Fortunately, the operator can rely on ultrasound system software to make these calculations, but measurements averaged from two orthogonal views (midesophageal four-chamber and two-chamber) are recommended, particularly when extensive wall-motion abnormalities are present. Reference values for LV systolic and diastolic volumes are provided in Table 4–3.
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LV mass is the total weight of the myocardium and is equal to the product of the volume of the myocardium and the specific density of cardiac muscle. LV mass can be derived from the transgastric mid-papillary short-axis view using a simple geometric cube formula:
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LV Mass = 0.8 × {1.04 × [(LVIDd + PWTd + SWTd)3 − (LVIDd)3]} + 0.6 g
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The formula is based on the assumption that the LV is a prolate ellipse and is accurate only when there are no major distortions in LV geometry. Since LV dimensions are cubed, even a small error in diameter measurements is significantly amplified. Calculation of LV mass by TEE is comparable with TTE; however, TEE measurements are higher by an average of 6 gm/m2 (see Table 4–2).12
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Indexing LV dimensions to body surface area remains controversial, because such indexing has been shown to underestimate the incidence of LV hypertrophy in overweight and obese individuals. Furthermore, the total LV mass differs between men and women even when indexed for body surface area, and it is unclear whether indexing improves the predictive value of these measures for the occurrence of cardiovascular events.
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Recently, three-dimensional (3D) echocardiography has been utilized to calculate LV dimensions and shapes. Since 3D echocardiography does not make assumptions about the shape of the ventricle, volumetric measurements have been shown to be more accurate and comparable to gold standards such as MRI. The 3D calculations of LV volumes and dimensions have been primarily obtained with the matrix 3D TTE transducer. It remains to be seen whether images acquired with the recently introduced 3D matrix TEE probe will demonstrate similar degrees of accuracy and reproducibility.
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Quantification of Right Ventricle (RV)
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The normal RV is a low-pressure and highly compliant cardiac chamber. While it typically appears smaller than the LV, in fact it has almost the same volume as the LV. Since it is crescent shaped, has greater trabeculations, and is structurally more complex, it is incompletely visualized from a single 2D echocardiographic view.13,14 When using TEE for RV assessment, the midesophageal four-chamber view, with the multiplane angle adjusted to 10° to 20° to maximize tricuspid annulus diameter, should be used for measurement of RV size, wall thickness, and fractional area change (Figure 4–18 and Table 4–4). Normal RV free wall thickness as measured in the midesophageal four-chamber view at end-diastole is generally less than 0.5 cm.12 The right ventricular outflow tract (RVOT) extends from the anterosuperior aspect of the RV to the pulmonary artery and includes the pulmonic valve (Figure 4–19). The diameter of the RVOT is best measured from the midesophageal RV inflow-outflow view (Figure 4–19).
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Quantification of Left Atrium (LA) and Right Atrium (RA)
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LA size should be measured at end-systole, because the LA achieves its greatest dimension at that point in the cardiac cycle.12 The LA gradually enlarges with worsening diastolic dysfunction, and therefore LA dimensions can be followed over time to assess response to therapy and to assess prognosis. TEE measurement of LA dimensions is problematic and unreliable. This is because in most cases the LA size exceeds the maximal span of the ultrasound beam, and hence cannot be measured accurately. It is recommended that the LA be estimated from multiple imaging planes. Also, for purposes of risk stratification, LA volumes assessed with transthoracic echo have shown a greater prognostic value than LA size. The reference ranges for LA diameter in men and women are 3.0 to 4.0 cm and 2.7 to 3.8 cm, respectively.12
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There are limited data available on the best transthoracic or TEE view to assess the RA size or volume. As with the LA, the RA size should be estimated from multiple echocardiographic windows. It is also believed that RA volume may be a more accurate and reproducible estimate of size than a linear measure.