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Principles of Echocardiographic Image Acquisition and Doppler Analysis
An understanding of the basic principles of ultrasound imaging and Doppler echocardiography is essential both during data acquisition and for correct interpretation of the ultrasound information. Although, at times, current instruments provide instantaneous images so clear and detailed that it seems as if we can “see” the heart and blood flow directly, in actuality we always are looking at images and flow data generated by complex analyses of ultrasound waves reflected and backscattered from the patient's body. Knowledge of the strengths, and more importantly the limitations, of this technique is critical for correct clinical diagnosis and patient management. On the one hand, echocardiography can be used for decision making with a high degree of accuracy in a variety of clinical settings. On the other hand, if an ultrasound artifact is mistaken for an anatomic abnormality, a patient could undergo other needless, expensive, and potentially risky diagnostic tests or therapeutic interventions.
In this chapter, a brief (and necessarily simplified) overview of the basic principles of cardiac ultrasound imaging and flow analysis is presented. The reader is referred to the Suggested Reading at the end of the chapter for more information on these subjects. Because the details of image processing, artifact formation, and Doppler physics become more meaningful with experience, some readers may choose to return to this chapter after reading other sections of this book and after participating in some echocardiographic examinations.
Ultrasound Waves
Sound waves are mechanical vibrations that induce alternate refraction and compression of any physical medium through which they pass ( Fig. 1.1 ). Like other waves, sound waves are described in terms of ( Table 1.1 ):
- ▪
Frequency: cycles per second, or hertz (Hz)
- ▪
Velocity of propagation
- ▪
Wavelength: millimeters (mm)
- ▪
Amplitude: decibels (dB)
TABLE 1.1
Ultrasound Waves
| Definition | Examples | Clinical Implications | |
|---|---|---|---|
| Frequency ( f ) | The number of cycles per second in an ultrasound wave f = cycles/s = Hz |
Transducer frequencies are measured in MHz (1,000,000 cycles/s). Doppler signal frequencies are measured in kHz (1000 cycles/s). |
Different transducer frequencies are used for specific clinical applications because the transmitted frequency affects ultrasound tissue penetration, image resolution, and the Doppler signal. |
| Velocity of propagation (c) | The speed that ultrasound travels through tissue | The average velocity of ultrasound in soft tissue is ≈1540 m/s. | The velocity of propagation is similar in different soft tissues (e.g., blood, myocardium, liver, fat) but is much lower in lung and much higher in bone. |
| Wavelength (λ) | The distance between ultrasound waves: λ = c/ f = 1.54 / f (MHz) |
Wavelength is shorter with a higher-frequency transducer and longer with a lower-frequency transducer. | Image resolution is greatest (≈1 mm) with a shorter wavelength (higher frequency). Depth of tissue penetration is greatest with a longer wavelength (lower frequency). |
| Amplitude (dB) | Height of the ultrasound wave or “loudness” measured in decibels (dB) | A log scale is used for dB. On the dB scale, 80 dB represents a 10,000-fold and 40 dB indicates a 100-fold increase in amplitude. |
A very wide range of amplitudes can be displayed using a gray-scale display for both imaging and spectral Doppler. |
Frequency (f) is the number of ultrasound waves in a 1-second interval. The units of measurement are hertz, abbreviated Hz, which simply means cycles per second. A frequency of 1000 cycles/s is 1 kilohertz (kHz), and 1 million cycles/s is 1 megahertz (MHz). Humans can hear sound waves with frequencies between 20 Hz and 20 kHz; frequencies higher than this range are termed ultrasound. Diagnostic medical ultrasound typically uses transducers with a frequency between 1 and 20 MHz.
The speed that a sound wave moves through the body, called the velocity of propagation ( c ), is different for each type of tissue. For example, the velocity of propagation in bone is much faster (about 3000 m/s) than in lung tissue (about 700 m/s). However, the velocity of propagation in soft tissues, including myocardium, valves, blood vessels, and blood, is relatively uniform, averaging about 1540 m/s.
Echo Math
Wavelength
Wavelength ( λ ) is the distance from peak to peak of an ultrasound wave. Wavelength can be calculated by dividing the frequency ( f in Hz) by the propagation velocity ( c in m/s):
Because the propagation velocity in the heart is constant at 1540 m/s, the wavelength for any transducer frequency can be calculated ( Fig. 1.2 ) as:
or
For example, the wavelength emitted by a 5-MHz transducer can be calculated as:
Wavelength is important in diagnostic applications for at least two reasons:
- ▪
Image resolution is no greater than 1 to 2 wavelengths (typically about 1 mm).
- ▪
The depth of penetration of the ultrasound wave into the body is directly related to wavelength; shorter wavelengths penetrate a shorter distance than longer wavelengths.
Thus there is an obvious trade-off between image resolution (shorter wavelength or higher frequency preferable) and depth penetration (longer wavelength or lower frequency preferable).
The acoustic pressure, or amplitude, of an ultrasound wave indicates the energy of the ultrasound signal. Power is the amount of energy per unit time. Intensity (I) is the amount of power per unit area:
This relationship shows that if ultrasound power is doubled, intensity is quadruped. Instead of using direct measures of pressure energy, ultrasound amplitude is described relative to a reference value using the decibel scale. Decibels (dB) are familiar to all of us as the standard description of the loudness of a sound.
Echo Math
Decibels
Decibels are logarithmic units based on a ratio of the measured amplitude (A 2 ) to a reference amplitude (A 1 ) such that:
Thus a ratio of 1000 to 1 is
a ratio of 100 to 1 is
and a ratio of 2 to 1 is
A simple rule to remember is that a 6-dB change represents a doubling or halving of the signal amplitude or that a 40-dB change represents a 100 times difference in amplitude ( Fig. 1.3 ).
If acoustic intensity is used instead of amplitude, the constant 10 replaces 20 in the equation so that a 3-dB change represents doubling, and a 20-dB change indicates a 100-fold difference in amplitude. Both these decibel scales are used to refer to transmitted or received ultrasound waves or to describe attenuation effects. The advantages of the decibel scale are that a very large range can be compressed into a smaller number of values and that low-amplitude (weak) signals can be displayed alongside very high-amplitude (strong) signals. In an echocardiographic image, amplitudes typically range from 1 to 120 dB. The decibel scale is the standard format both for echocardiographic image display and for the Doppler spectral display, although other amplitude scales are sometimes available.
Ultrasound Tissue Interaction
Propagation of ultrasound waves in the body to generate ultrasound images and Doppler data depends on a tissue property called acoustic impedance ( Table 1.2 ). Acoustic impedance ( Z) depends on tissue density ( ρ ) and on the propagation velocity in that tissue ( c ):
TABLE 1.2
Ultrasound Tissue Interaction
| Definition | Examples | Clinical Implications | |
|---|---|---|---|
| Acoustic impedance ( Z ) | A characteristic of each tissue defined by tissue density (ρ) and propagation of velocity (c) as: Z = ρ × c |
Lung has a low density and slow propagation velocity, whereas bone has a high density and fast propagation velocity. Soft tissues have smaller differences in tissue density and acoustic impedance. |
Ultrasound is reflected from boundaries between tissues with differences in acoustic impedance (e.g., blood vs. myocardium). |
| Reflection | Return of ultrasound signal to the transducer from a smooth tissue boundary | Reflection is used to generate 2D cardiac images. | Reflection is greatest with the ultrasound beam in perpendicular to the tissue interface. |
| Scattering | Radiation of ultrasound in multiple directions from a small structure (e.g., blood cells) | The change in frequency of signals scattered from moving blood cells is the basis of Doppler ultrasound. | The amplitude of scattered signals is 100 to 1000 times less than reflected signals. |
| Refraction | Deflection of ultrasound waves from a straight path due to differences in acoustic impedance | Refraction is used in transducer design to focus the ultrasound beam. | Refraction in tissues results in double image artifacts. |
| Attenuation | Loss in signal strength due to absorption of ultrasound energy by tissues | Attenuation is frequency dependent with greater attenuation (less penetration) at higher frequencies. | A lower-frequency transducer is needed for apical views or in larger patients on transthoracic imaging. |
| Resolution | The smallest resolvable distance between two specular reflectors on an ultrasound image | Resolution has three dimensions: along the length of the beam (axial), lateral across the image (azimuthal), and in the elevational plane. | Axial resolution is most precise (as small as 1 mm), so imaging measurements are best made along the length of the ultrasound beam. |
Although the velocity of propagation differs among tissues, tissue density is the primary determinant of acoustic impedance for diagnostic ultrasound. Lung tissue has a very low density compared with bone, which has a very high density. Soft tissues, such as blood and myocardium, have much smaller differences in tissue density and acoustic impedance. Acoustic impedance determines the transmission of ultrasound waves through a tissue; differences in acoustic impedance result in reflection of ultrasound waves at tissue boundaries.
The interaction of ultrasound waves with the organs and tissues of the body can be described in terms of ( Fig. 1.4 ):
- ▪
Reflection
- ▪
Scattering
- ▪
Refraction
- ▪
Attenuation
Reflection
The basis of ultrasound imaging is reflection of the transmitted ultrasound signal from internal structures. Ultrasound is reflected at tissue boundaries and interfaces, with the amount of ultrasound reflected dependent on the:
- ▪
Difference in acoustic impedance between the two tissues
- ▪
Angle of reflection
Smooth tissue boundaries with a lateral dimension greater than the wavelength of the ultrasound beam act as specular, or “mirror-like,” reflectors. The amount of ultrasound reflected is constant for a given interface, although the amount received back at the transducer varies with angle because (like light reflected from a mirror) the angle of incidence and reflection is equal. Thus optimal return of reflected ultrasound occurs at a perpendicular angle (90°). Remembering this fact is crucial for obtaining diagnostic ultrasound images. It also accounts for ultrasound “dropout” in a two-dimensional (2D) or three-dimensional (3D) image when too little or no reflected ultrasound reaches the transducer resulting from a parallel alignment between the ultrasound beam and tissue interface.
Scattering
Scattering of the ultrasound signal, instead of reflection, occurs with small structures, such as red blood cells suspended in fluid, because the radius of the cell (about 4 µm) is smaller than the wavelength of the ultrasound signal. Unlike a reflected beam, scattered ultrasound energy radiates in all directions. Only a small amount of the scattered signal reaches the receiving transducer, and the amplitude of a scattered signal is 100 to 1000 times (40–60 dB) less than the amplitude of the returned signal from a specular reflector. Scattering of ultrasound from moving blood cells is the basis of Doppler echocardiography.
The extent of scattering depends on:
- ▪
Particle size (red blood cells)
- ▪
Number of particles (hematocrit)
- ▪
Ultrasound transducer frequency
- ▪
Compressibility of blood cells and plasma
Although experimental studies show differences in backscattering with changes in hematocrit, variation over the clinical range has little effect on the Doppler signal. Similarly, the size of red blood cells and the compressibility of blood cells and plasma do not change significantly. Thus the primary determinant of scattering is transducer frequency.
Scattering also occurs within tissues, such as the myocardium, from interference of backscattered signals from tissue interfaces smaller than the ultrasound wavelength. Tissue scattering results in a pattern of speckles; tissue motion can be measured by tracking these speckles from frame to frame, as discussed in Chapter 4 .
Refraction
Ultrasound waves can be refracted —deflected from a straight path—as they pass through a medium with a different acoustic impedance. Refraction of an ultrasound beam is analogous to refraction of light waves as they pass through a curved glass lens (e.g., prescription eyeglasses). Refraction allows enhanced image quality by using acoustic “lenses” to focus the ultrasound beam. However, refraction also occurs in unplanned ways during image formation, with resulting ultrasound artifacts, most notably the “double-image” artifact.
Attenuation
Attenuation is the loss of signal strength as ultrasound interacts with tissue. As ultrasound penetrates into the body, signal strength is progressively attenuated because of absorption of the ultrasound energy by conversion to heat, as well as by reflection and scattering. The degree of attenuation is related to several factors, including the:
- ▪
Attenuation coefficient of the tissue
- ▪
Transducer frequency
- ▪
Distance from the transducer
- ▪
Ultrasound intensity (or power)
The attenuation coefficient (α) for each tissue is related the decrease in ultrasound intensity (measured in dB) from one point (I 1 ) to a second point (I 2 ) separated by a distance (l) as described by the equation:
The attenuation coefficient for air is very high (about 1000×) compared with soft tissue so that any air between the transducer and heart results in substantial signal attenuation. This is avoided on transthoracic examinations by use of a water-soluble gel to form an airless contact between the transducer and the skin; on transesophageal echocardiography (TEE) examination, attenuation is avoided by maintaining close contact between the transducer and esophageal wall. The air-filled lungs are avoided by careful patient positioning and the use of acoustic “windows” that allow access of the ultrasound beam to the cardiac structures without intervening lung tissue. Other intrathoracic air (e.g., pneumomediastinum, residual air after cardiac surgery) also results in poor ultrasound tissue penetration because of attenuation, resulting in suboptimal image quality.
The power output of the transducer is directly related to the overall degree of attenuation. However, an increase in power output causes thermal and mechanical bioeffects as discussed in “Bioeffects and Safety,” p. 27.
Overall attenuation is frequency dependent such that lower ultrasound frequencies penetrate deeper into the body than higher frequencies. The depth of penetration for adequate imaging tends to be limited to approximately 200 wavelengths. This translates roughly into a penetration depth of 30 cm for a 1-MHz transducer, 6 cm for a 5-MHz transducer, and 1.5 cm for a 20-MHz transducer, although diagnostic images at depths greater than these postulated limits can be obtained with state-of-the-art equipment. Thus attenuation, as much as resolution, dictates the need for a particular transducer frequency in a specific clinical setting. For example, visualization of distal structures from the apical approach in a large adult patient often requires a low-frequency transducer. From a TEE approach, the same structures can be imaged (at better resolution) with a higher-frequency transducer. The effects of attenuation are minimized on displayed images by using different gain settings at each depth, an instrument control called time-gain (or depth-gain) compensation.
Transducers
Piezoelectric Crystal
Ultrasound transducers use a piezoelectric crystal both to generate and to receive ultrasound waves ( Fig. 1.5 ). A piezoelectric crystal is a material (e.g., quartz or a titanate ceramic) with the property that an applied electric current results in alignment of polarized particles perpendicular to the face of the crystal with consequent expansion of crystal size. When an alternating electric current is applied, the crystal alternately compresses and expands, generating an ultrasound wave. The frequency that a transducer emits depends on the nature and thickness of the piezoelectric material.
Conversely, when an ultrasound wave strikes the piezoelectric crystal, an electric current is generated. Thus the crystal can serve both as a “receiver” and as a “transmitter.” Basically, the ultrasound transducer transmits a brief burst of ultrasound and then switches to the “receive mode” to await the reflected ultrasound signals from the intracardiac acoustic interfaces. This cycle is repeated temporally and spatially to generate ultrasound images. Image formation is based on the time delay between ultrasound transmission and return of the reflected signal. Deeper structures have a longer time of flight than shallower structures, with the exact depth calculated based on the speed of sound in blood and the time interval between the transmitted burst of ultrasound and return of the reflected signal.
The burst, or pulse, of ultrasound generated by the piezoelectric crystal is very brief, typically 1 to 6 µs, because a short pulse length results in improved axial (along the length of the beam) resolution. Damping material is used to control the ring-down time of the crystal and hence the pulse length. Pulse length also is determined by frequency because a shorter time is needed for the same number of cycles at higher frequencies. The number of ultrasound pulses per second is called the pulse repetition frequency (PRF). The total time interval from pulse to pulse is called the cycle length, with the percentage of the cycle length used for ultrasound transmission called the duty factor . Ultrasound imaging has a duty factor of about 1% compared with 5% for pulsed Doppler and 100% for continuous-wave (CW) Doppler. The duty factor is a key element in the patient's total ultrasound exposure.
The range of frequencies contained in the pulse is described as its frequency bandwidth. A wider bandwidth allows better axial resolution because of the ability of the system to produce a narrow pulse. Transducer bandwidth also affects the range of frequencies that can be detected by the system with a wider bandwidth, which allows better resolution of structures distant from the transducer. The stated frequency of a transducer represents the center frequency of the pulse.
Types of Transducers
The simplest type of ultrasound transducer is based on a single piezoelectric crystal ( Table 1.3 ). Alternate pulsed transmission and reception periods allow repeated sampling along a single line, with the sampling rate limited only by the time delay needed for return of the reflected ultrasound wave from the depth of interest. An example of using the transducer for simple transmission and reception along a single line is an A-mode (amplitude vs. depth) or M-mode (depth vs. time) cardiac recording when a high sampling rate is desirable.
TABLE 1.3
Ultrasound Transducers
| Definition | Examples | Clinical Implications | |
|---|---|---|---|
| Type | Transducer characteristics and configuration Most cardiac transducers use a phased array of piezoelectric crystals. |
TTE (adult and pediatric) Nonimaging CW Doppler 3D echocardiography TEE Intracardiac |
Each transducer type is optimized for a specific clinical application. More than one transducer may be needed for a full examination. |
| Transmission frequency | The central frequency emitted by the transducer | Transducer frequencies vary from 2.5 MHz for TTE to 20 MHz for intravascular imaging. | A higher-frequency transducer provides improved resolution but less penetration. Doppler signals are optimal at a lower transducer frequency than used for imaging. |
| Power output | The amount of ultrasound energy emitted by the transducer | An increase in transmitted power increases the amplitude of the reflected ultrasound signals. | Excessive power output may result in bioeffects measured by the mechanical and thermal indexes. |
| Bandwidth | The range of frequencies in the ultrasound pulse | Bandwidth is determined by transducer design. | A wider bandwidth allows improved axial resolution for structure distant from the transducer. |
| Pulse (or burst) length | The length of the transmitted ultrasound signal | A higher-frequency signal can be transmitted in a shorted pulse length compared with a lower-frequency signal. | A shorter pulse length improves axial resolution. |
| Pulse repetition frequency (PRF) | The number of transmission-receive cycles per second | The PRF decreases as imaging (or Doppler) depth increases because of the time needed for the signal to travel from and to the transducer. | Pulse repetition frequency affects image resolution and frame rate (particularly with color Doppler). |
| Duty factor | The percentage to time that ultrasound is transmitted | It ranges from ≈1% for 2D imaging, to 5% for pulsed Doppler, and 100% for CW Doppler. | A higher duty factor means more tissue exposure to ultrasound. |
| Focal depth | Beam shape and focusing are used to optimize ultrasound resolutions at a specific distance from the transducer | Structures close to the transducer are best visualized with a short focal depth, and distant structures are best visualized with a long focal depth. | The length and site of a transducer's focal zone are primarily determined by transducer design but are adjustable during the examination. |
| Aperture | The surface of the transducer face where ultrasound is transmitted and received | A small nonimaging CW Doppler transducer allows optimal positioning and angulation of the ultrasound beam. | A larger aperture allows a more focused beam A smaller aperture allows improved transducer angulation on TTE imaging. |
Formation of more complex images uses an array of ultrasound crystals arranged to provide a 2D tomographic or 3D volumetric data set of signals. Each element in the transducer array can be controlled electronically both to direct the ultrasound beam across the region of interest and to focus the transmitted and received signals. Echocardiographic imaging uses a sector scanning format with the ultrasound signal originating from a single location (the narrow end of the sector), thus resulting in a fanlike shape of the image. Sector scanning is optimal for cardiac applications because it allows a fast frame rate to show cardiac motion and a small transducer size (aperture or “footprint”) to fit into the narrow acoustic windows used for echocardiography. 3D imaging transducers are discussed in Chapter 4 .
Most transducers can provide simultaneous imaging and Doppler analysis, for example, 2D imaging and a superimposed color Doppler display. Quantitative Doppler velocity data are recorded with the image “frozen” or with only intermittent image updates, with the ultrasound crystals used to optimize the Doppler signal. Although CW Doppler signals can be obtained using two elements of combined transducer, use of a dedicated nonimaging transducer with two separate crystals (with one crystal continuously transmitting and the other continuously receiving the ultrasound waves) is recommended when accurate high-velocity recordings are needed. The final configuration of a transducer depends on transducer frequency (higher-frequency transducers are smaller) and beam focusing, as well as the intended clinical use, for example, transthoracic versus TEE imaging.
Beam Shape and Focusing
An unfocused ultrasound beam is shaped like the light from a flashlight, with a tubular beam for a short distance that then diverges into a broad cone of light ( Fig. 1.6 ). Even with current focused transducers, ultrasound beams have a 3D shape that affects measurement accuracy and contributes to imaging artifacts. Beam shape and size depend on several factors, including:
- ▪
Transducer frequency
- ▪
Distance from the transducer
- ▪
Aperture size and shape
- ▪
Beam focusing
Aperture size and shape and beam focusing can be manipulated in the design of the transducer, but the effects of frequency and depth are inherent to ultrasound physics. For an unfocused beam, the initial segment of the beam is columnar in shape (near-field F n ) with a length dependent on the diameter (D) of the transducer face and wavelength (λ):
For a 3.5-MHz transducer with a 5-mm diameter aperture, this corresponds to a columnar length of 1.4 cm. Beyond this region, the ultrasound beam diverges (far field), with the angle of divergence θ determined as:
This equation indicates a divergence angle of 6° beyond the near field, resulting in an ultrasound beam width of about 4.4 cm at a depth of 20 cm for this 3.5-MHz transducer. With a 10-mm diameter aperture, F n would be 5.7 cm, and beam width at 20 cm would be about 2.5 cm ( Fig. 1.7 ).
The shape and focal depth (narrowest point) of the primary beam can be altered by making the surface of the piezoelectric crystal concave or by the addition of an acoustic lens. This allows generation of a beam with optimal characteristics at the depth of most cardiac structures, but again, divergence of the beam beyond the focal zone occurs. Some transducers allow manipulation of the focal zone during the examination. Even with focusing, the ultrasound beam generated by each transducer has a lateral and an elevational dimension that depends on the transducer aperture, frequency, and focusing. Beam geometry for phased-array transducers also depends on the size, spacing, and arrangement of the piezoelectric crystals in the array.
In addition to the main ultrasound beam, dispersion of ultrasound energy laterally from a single-crystal transducer results in formation of side lobes at an angle θ from the central beam where sin θ = mλ/ D, and m is an integer describing sequential side lobes (i.e., 1, 2, 3, and so on) ( Fig. 1.8 ). Reflected or backscattered signals from these side lobes are received by the transducer, resulting in image or flow artifacts. With phased-array transducers, additional accessory beams at an even greater angle from the primary beam, termed grating lobes, also occur as a result of constructive interference of ultrasound wave fronts. Both the side lobes and the grating lobes affect the lateral and elevational resolution of the transducer.
Resolution
Image resolution occurs for each of three dimensions ( Fig. 1.9 ):
- ▪
A xial resolution along the length of the ultrasound beam
- ▪
Lateral resolution side to side across the 2D image
- ▪
Elevational resolution or thickness of the tomographic slice
Of these three, axial resolution is most precise, so quantitative measurements are made most reliably using data derived from a perpendicular alignment between the ultrasound beam and structure of interest. Axial resolution depends on the transducer frequency, bandwidth, and pulse length but is independent of depth ( Table 1.4 ). Determination of the smallest resolvable distance between two specular reflectors with ultrasound is complex but is typically about twice the transmitted wavelength; higher-frequency (shorter-wavelength) transducers have greater axial resolution. For example, with a 3.5-MHz transducer, axial resolution is about 1 mm, versus 0.5 mm with a 7.5-MHz transducer. A wider bandwidth also improves resolution by allowing a shorter pulse, thus avoiding overlap between the reflected ultrasound signals from two adjacent reflectors.
TABLE 1.4
Determinants of Resolution in Ultrasound Imaging
| Axial Resolution |
|
| Lateral Resolution |
|
| Elevational Resolution |
|
Lateral resolution varies with the depth of the specular reflector from the transducer, primarily related to beam width at each depth. In the focal region where beam width is narrow, lateral resolution approaches axial resolution, and a point target will appear as a point on the 2D image. At greater depths, beam width diverges so a point target results in a reflected signal as wide as the width of the beam and accounts for “blurring” of images in the far field. If the 2D image is examined carefully, progressive widening of the echo signals from similar targets along the length of the ultrasound beam can be appreciated ( Fig. 1.10 ). Erroneous interpretations occur when the effects of beam width are not recognized. For example, beam width artifact from a strong specular reflector often appears to be an abnormal linear structure. Other factors that affect lateral resolution are transducer frequency, aperture, bandwidth, and side and grating lobe levels.
Resolution in the elevational plane is more difficult to recognize on the 2D image but is equally important in clinical diagnosis. The thickness of the tomographic plane varies across the 2D image, depending on transducer design and focusing, both of which affect beam width in the elevational plane at each depth. In general, cardiac ultrasound images have a “thickness” of approximately 3 to 10 mm depending on depth and the specific transducer used. The tomographic image generated by the instrument, in effect, includes reflected and backscattered signals from this entire thickness. Strong reflectors adjacent to the image plane can appear to be “in” the image plane because of elevational beam width. Even more distant strong reflectors sometimes appear superimposed on the tomographic plane because of side lobes in the elevational plane. For example, a linear echo in the aortic lumen from an adjacent calcified atheroma may look like a dissection flap. These principles of ultrasound imaging also apply to 3D echocardiography.
Ultrasound Imaging Modalities
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