Basics of ultrasound

Ultrasound waves can provide tissue information through reflectioned at the boundary between different acoustic media. The subsequently generated echoes are sent back to the transducer, which is then used for image composition

The human ear can detect sound waves with a frequency between 20 Hz to 20,000 Hz. Acoustic signals with a higher frequency (=ultrasound) physically behave just as audible sound waves. For medical applications, ultrasound is used with a frequency in the range of MHz.

Sound waves are mechanical vibrations that can be described in terms of frequency or Hertz (Hz), ie, the number of repetitions or cycles per second. Other characteristics include wavelength, the distance between excitations, measured in mm; and the amplitude of excitation, measured in decibels (dB). A 6-dB change results in a doubling (or halving) of the signal amplitude. Medical ultrasound imaging typically uses sound waves at frequencies of 1,000,000 to 20,000,000 Hz (1.0 to 20 MHz). Frequency and wavelength are mathematically related to the velocity of the ultrasound beam within the tissue (approximately 1,5400,000 mm/sec for human tissue) as indicated by the following equations:

Propagation velocity = Wavelength (mm) x frequency (Hz)

Wavelength (mm) = 1,540,000 mm/sec / frequency (Hz)

Wavelength (mm) = 1.54 / frequency (MHz)

The resolution of a recording, ie, the ability to distinguish two objects that are spatially close together, varies directly with the frequency and inversely with the wavelength. High frequency, short wavelength ultrasound can separate objects that are less than 1 mm apart. Echocardiographic image resolution is generally 1 or 2 wavelengths. Thus, imaging with a 2.5-MHz transducer would result in a resolution of approximately 1 mm. Imaging with higher frequency (and lower wavelength) transducers permits enhanced spatial resolution. However, because of attenuation, the depth of tissue penetration or the ability to transmit sufficient ultrasonic energy into the chest is directly related to wavelength and therefore inversely related to transducer frequency. As a result, the trade-off for use of higher frequency transducers is reduced tissue penetration. The trade-off between tissue resolution and penetration guides the choice of transducer frequency for clinical imaging. As an example, higher frequency transducers can be used in echocardiography for imaging of structures close to the transducer or the chest wall, such as the apex of the left ventricle with transthoracic imaging. When an ultrasonic beam travels through a homogeneous medium, its path is a straight line. However, when the medium is not homogeneous or when the beam travels through a medium with two or more interfaces, its path is altered. The relationship between ultrasound waves and tissues can be described in terms of reflection, scattering, refraction, and attenuation. The last three factors all act to decrease the magnitude of the ultrasound wave. The potential of diagnostic ultrasound to display structures and tissue is influenced by a number of factors related to the way ultrasound waves behave in tissue. In addition, ultrasound waves undergo attenuation and therefore have a limited depth of penetration. Most importantly, ultrasound can create artifacts which may significantly alter image quality and falsely display structures or tissue. Various examples of ultrasound artifacts include attenuation, acoustic shadowing, near field clutter, reverberations, mirror artifacts, side lobes, beam width artifacts, stitching artifacts in 3D imaging and erroneous machine settings. An overview of a number of possible artifacts is provided in the below section.

When an ultrasound beam “hits” a tissue boundary/interface, a certain amount of the ultrasound is reflected back to the transducer, like a mirror. The magnitude of the reflected wave is dependent on the acoustic impedance of the tissue:

Acoustic impedance = tissue density x propagation velocity

Tissues with increased density reflect a greater proportion of the ultrasound beam. The magnitude of the reflected beam which is received by the transducer is dependent upon the angle between the ultrasound beam and tissue interface. Since the angle of incidence equals the angle of reflection, the “optimal” return of the reflected ultrasound occurs at a 90? (perpendicular) orientation.

Small structures, eg, less than 1 wavelength in lateral dimension, result in scattering of the ultrasound signal. Unlike a reflected beam, scattering results in the ultrasound beam being radiated in all directions, with minimal signal returning to the transducer.

Ultrasound waves can all be refracted, or deflected from their orientation, as they pass into a medium of different acoustic impedance.

The ultrasound signal strength is progressively reduced due to absorption of the ultrasound energy by conversion to heat, a process called attenuation. Attenuation is frequency and, from the above formation, wavelength dependent. The depth of penetration is limited to approximately 200 wavelengths, corresponding to a depth of 30 cm for a 1 MHz transducer, 12 cm for 2.5 MHz transducer, and 6 cm for a 5 MHz transducer. Attenuation is also dependent upon acoustic impedance and any mismatch in impedance between adjacent structures. Air has a very high acoustic impedance, resulting in significant signal attenuation when imaging through lung tissue, especially emphysematous lung, or pathologic conditions such as pneumomediastinum or subcutaneous emphysema. In contrast, filling of the pleural space with fluid, generally enhances ultrasound imaging.

Ultrasound transducers use piezoelectric crystals to both generate and receive ultrasound waves. These crystals (quartz or titanate ceramic) alternately compress and expand the alternating electric current that is applied, thereby generating the ultrasound wave. Following a brief period of transmission, typically 1 to 6 microseconds, the same crystal also acts as a receiver. When a reflected ultrasound wave impacts the piezoelectric crystal, an electric current is generated. Image formation, which is related to the distance of a structure from the transducer, is based upon the time interval between ultrasound transmission and arrival of the reflected signal. The amplitude is proportional to the incident angle and acoustic impedance, and timing is proportional to the distance from the transducer. The simplest type of ultrasound transducer has a single piezoelectric crystal and is often used for M-mode recordings. Generation of a 2D image requires mechanical or electronic “sweeping” of the ultrasound beam across the plane of interest or sector. Initially, mechanical transducers physically moved a crystal. Today, phased-array transducers consist of a series of ultrasound crystals arranged so that they can be “electronically” steered, with no moving parts. The phased-array transducers are the most common type currently used for clinical echocardiography. In contrast to echocardiographic imaging, continuous-wave Doppler examinations utilize a pair of dedicated crystals: one for continuous transmission; and one for continuous receiving.

If the propagation velocity of the ultrasonic waves is fixed (the time between the transmission of an ultrasound and receiving the echo) it corresponds to a particular distance traveled. Thus, if the speed of the sound is fixed, then the depth of the reflecting interface is proportional to the time. In other words the later the echo, the deeper the reflective interface. Traveling speed depends on the intermediate substance. In order to produce images, it is assumed that the propagation velocity amounts 1540 m/s in soft tissue.

In the most simple (historical) Amplitude-mode, the strength of the echo is picked up on a scan line plotted as a function of time (=depth). The A-mode echo strength is compensated for attenuation (ie the depth or time).

Image resolution with 2D echocardiography can be considered in terms of:

“Axial” resolution along the length of the ultrasound beam

“Lateral” resolution

“Elevational” resolution which is the thickness of the tomographic “slice”

Axial resolution is a function of the transducer frequency, bandwidth, and pulse length. Since the smallest resolvable distance between two specular reflections is 1 wavelength, higher-frequency transducers result in enhanced axial resolution. A wider bandwidth also improves resolution by allowing for a shorter pulse. Lateral resolution varies with transducer frequency, beam width, bandwidth, aperture (width) of the transducer, and side lobes. At greater depths, beam width diverges so that a point target results in a reflected signal as wide as the beam width. Beam width artifacts appear as a bright linear structure. The 2D tomographic image includes reflected and backscattered signals from the entire thickness. The thickness of the 2D image is variable over the image plane and is dependent upon the transducer design and focusing. Most clinical images have a “thickness” of 3 to 10 mm, depending on depth. Strong reflectors near the image plane may appear “in” the image plane due to elevational beam width.

B-mode imaging, or brightness modulation, displays the varying intensities of the returning echoes as varying degrees of brightness, instead of the vertical deflections used in A-mode. The brightness of the amplitudes in B-mode is represented as pixels. Echoes with greater intensity are displayed with greater degrees of brightness. Multiple scan lines are combined to one image. These scan lines are sequentially obtained by to steer the ultrasound beam in different directions.

Motion or “M”-mode echocardiography is among the earliest forms of cardiac ultrasound. With this technique, a single crystal rapidly alternates between transmission and receiver modes with rapid updating (>1000 Hz); as a result, rapidly moving structures (eg, valve leaflets) can be monitored for their characteristic motion. M-mode data can be recorded on paper or displayed on the video monitor at sweep speeds of 50 to 100 mm/sec. Although originally performed using dedicated crystals, alignment of the M-mode beam is now typically performed with 2D imaging guidance.

A moving target will backscatter an ultrasound beam to the transducer so that the frequency observed when the target is moving toward the transducer is higher and the frequency observed when the target is moving away from the transducer is lower than the original transmitter frequency. This Doppler phenomenon is familiar to us as the sound of a train whistle as it moves toward (higher frequency) or away (lower frequency) from the observer. This difference in frequency between the transmitted frequency (F[t]) and received frequency (F[r]) is the Doppler shift: Doppler shift (F[d]) = F[r] – F[t] Blood flow velocity (V) is related to the Doppler shift by the speed of sound in blood (C) and α, the intercept angle between the ultrasound beam and the direction of blood flow. A factor of 2 is used to correct for the “round-trip” transit time to and from the transducer.

F[d] = 2 x F[t] x [(V x cos α)] / C

This equation can be solved for V, by substituting (F[r] – F[t]) for F[d]: V = [(F[r] -F[t]) x C] / (2 x F[t] x cos α). Note that the angle of the ultrasound beam and the direction of blood flow are critically important in the calculation.

Continuous wave Doppler employs two dedicated ultrasound crystals: one for continuous transmission and a second for continuous reception of ultrasound signals. This permits measurement of very high frequency Doppler shifts or velocities. The “cost” is that this technique receives a continuous signal along the entire length of the ultrasound beam. Thus, there may be overlap in certain settings, such as stenoses in series (eg, left ventricular outflow tract gradient and aortic stenosis) or flows that are in close proximity/alignment (eg, aortic stenosis and mitral regurgitation). Differentiation of the signal from each component may still be determined from the characteristic timing and/or profile. An ideal Doppler profile is one with a smooth “outer” contour, well-defined edge and maximum velocity, and abrupt onset and termination. The continuous wave Doppler profile is usually “filled in” because lower-velocity signals proximal and distal to the point of maximum velocity are also recorded. Although the maximum frequency shift depends on α, the profile, onset, and termination of the Doppler signal are not dependent upon this value, resulting in inappropriate underestimation of true velocity. For this reason, continuous wave Doppler positioning is often integrated with two-dimensional (2D) and color flow imaging to allow for good alignment with flow (ie, angle less than 20°). Continuous wave Doppler is typically used to measure higher velocities as in pulmonary hypertension and aortic stenosis.

In contrast to continuous wave Doppler, which records signal along the entire length of the ultrasound beam, pulsed wave Doppler permits sampling of local blood flow velocities about a specific region. This modality is particularly useful for assessing the relatively low velocity flows associated with transmitral or transtricuspid blood flow, pulmonary venous flow, left atrial appendage flow, or for confirming the location of eccentric jets of aortic regurgitation and mitral regurgitation. To permit this, a pulse of ultrasound is transmitted and then the receiver “listens” during a subsequent interval defined by the distance from the transmitter and the sample site. This transducer mode of transmit-wait-receive is repeated at an interval termed the pulse-repetition frequency (PRF). The PRF is therefore depth-dependent, being greater for near regions and lower for distant or deeper regions. The distance from the transmitter to the region of interest is called the sample volume, with the width and length of the sample volume varied by adjusting the length of the transducer “receive” interval. In contrast to continuous wave Doppler, which is sometimes performed without 2D guidance, pulsed Doppler is always performed with 2D guidance to determine the sample volume position Because pulsed wave Doppler echo repeatedly samples the returning signal, there is a maximum limit to the frequency shift or velocity that can be measured unambiguously. Correct identification of the frequency of an ultrasound waveform requires sampling at least twice per wavelength. Thus, the maximum detectable frequency shift or the Nyquist limit is one-half the PRF. If the velocity of interest exceeds the Nyquist limit, “wraparound” of the signal occurs first into the reverse channel, then back to the forward channel; this is known as aliasing. Techniques that can minimize aliasing during pulsed Doppler include using a lower frequency transducer and shifting the baseline. Another solution is to increase the number of sample volumes, or high PRF. As noted above, when a pulse is transmitted, backscatter along the entire length of the beam is received. Depth resolution is achieved with pulsed Doppler using the duration of the “wait” period. However, signals from exactly twice (or 3x, 4x, etc) the distance will reach the transducer during the “receive” phase of the next (or subsequent) cycle. As a result, signals from 1x, 2x, 3x, 4x, 5x, etc have the potential for confounding the analysis. The latter signals are generally of low amplitude and do not interfere with the spectral display. If, however, the sample volume is deliberately placed at one-half the depth of interest, backscattered signals from the 2x sample volume, the true depth of interest, will return to the transducer during the “receive” phase of the following cycle. This recording of signal at a higher PRF permits measurement of higher velocities without signal averaging. Even greater velocities could be achieved using additional sample volumes. In the pulsed Doppler multiple short ultrasound pulses are transmitted by a transmitter and the echo is recorded by the same transmitter. Comparison of the emitted with the signals received supplies again the Doppler shift, and therefore the velocity of the blood. The advantage of the technique is that the Doppler shift can be measured at a pre-selected depth. Which, after all, is equivalent to echoes at a given time after the transmission of the pulse. The disadvantage of the art is that no high blood velocities can be measured. This technique can be combined with B-mode; This is called duplex.

The velocity of blood flow are usually shown as a graph in function of time. Positive blood flow velocity away from the transducer, negative blood flow velocity towards the transducer. Due to the fact that high speeds lead to aliasing in Pulsed wave Doppler, these measurements are not reliable with this technique and are to be measured with Continuous wave Doppler.

Color flow imaging is a form of pulsed wave Doppler. A scan line uses several short successive pulse-echo sequences. Moving structures in the scanline are to cause a phase shift This will bedisplayed on the image in color. These measurements are less accurate than when measured with pulsed wave Doppler. However, the advantage of Color Doppler is that information from various depths are obtained immediately. . Pulse 1 reaches the moving reflector. Pulse 2, emitted shortly after pulse 1, will reach the moving reflector at a different point, as the moving reflector in the meantime has travelled a distance d. By the time the echo of the second pulse reaches the reference point, the echo from pulse 1 has travelled a distance of 2 x d (for the surface). By comparison of the two echoes and the resulted phase shift due to the moving reflector, the Doppler shift can be derived. Flow towards the ultrasound transducer appears in red. Flow away, from the transducer appears in blue.

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