Fundamentals: Essential technology, concepts, and capability

ASHOT E. SARGSYAN, MICHAEL BLAIVAS, PHILIP LUMB and DIMITRIOS KARAKITSOS

“There is geometry in the humming of the strings . . . there is music in the spacing of the spheres”

Christian Doppler was born in Salzburg, Austria on November 29, 1803 and lived a short and deprived life, like many scientists of his time. In an 1842 session of the Science Section of the Royal Bohemian Society in Prague, he presented a thesis entitled “Concerning the colored light of double stars and other celestial constellations.” Other milestones relevant to this chapter include the discovery of piezoelectric phenomenon (Curie brothers, 1880); the construction of the first sonar (Langevin-Chilowski, 1916); and the early efforts to use ultrasound for diagnostic purposes (Karl Dussik, 1942), which, along with the technologic progress in the second half of the twentieth century, paved the way to modern ultrasound imaging. Despite tremendous advances in ultrasound technology over the past 60 years, its basic principles are still the same: operation of piezoelectric sonar with frequency analysis capability.

Fundamentals: Principles, terms, and concepts

Ultrasound is a mechanical wave that requires a medium to travel (i.e., human tissue), with a frequency above the audible range ceiling of 20 kHz. Ultrasound systems are tomographic devices that transmit short pulses of ultrasound into the body and measure the round-trip time and intensity of each of the numerous echoes returning after the pulse. The time of arrival of an echo determines the distance from the transducer, that is, the location of its source in the body. The intensity of the echo is converted to brightness of a given point in the image. In other words, each pixel (element of the image) on the display device corresponds to a point inside the body, and its brightness depends on the strength of the echo that came from that location. Together, all pixels form a grayscale tomographic image. Parts of the image with mostly bright pixels (a brighter overall appearance) are termed hyperechoic, as opposed to hypoechoic (darker) areas. The relative ability of an organ or tissue to produce echoes is called echogenicity, that is, tissues or structures producing hyperechoic image are considered more echogenic.1–7 Parts of the image with only black pixels are called anechoic or echo-free and mostly correspond to homogenous liquids (e.g., blood, urine, effusion, cystic fluid).

Frequency (measured in cycles per second [hertz, Hz]) is the number of wave cycles in 1 second. Frequency is determined solely by the sound source and not by the medium. Frequencies used by general-purpose ultrasound machines range between 2 and 15 megahertz (MHz). Higher frequencies, up to 40 MHz, are used for intravascular and other catheter-based applications and in specialized ophthalmologic and dermatologic techniques. Propagation speed is the velocity of sound in a given medium and is determined solely by the characteristics of the medium, such as density and stiffness (does not depend on the source of sound or its frequency). Ultrasound travels through soft tissues at a speed of approximately 1.54 mm/μsec, or 1540 m/sec). The stiffer the tissue, the greater the propagation speed (Figure 1-1). Ultrasound waves are generated by piezoelectric crystals (e.g., lead zirconate titanate, or PZT) that convert electrical energy into mechanical energy and vice versa (see Figure 1-1). Electrical pulses or short bursts of alternating voltage stimulate crystals to produce ultrasound pulses in the medium, causing displacement and oscillation of its molecules. Pressure change–Velocity of such oscillations in response to sound pressure determines the acoustic impedance (lower velocities correspond to higher impedance). As ultrasound passes from one medium to another (e.g., from gas to liquid), an impedance gradient at the tissue boundary causes a part of the energy to form a reflected wave (echo) while the remainder of the energy proceeds into the second medium.^1–7 Reflection occurs every time the ultrasound pulse encounters a new boundary (reflector). Specular (mirror-like) reflectors are smooth and flat boundaries larger than the pulse dimensions (e.g., diaphragm, walls of a major vessels). The echo reflection angle equals the angle of incidence; when the beam strikes a specular reflector at 90 degrees (normal incidence, Figure 1-2), a very strong echo travels back toward the source. Nonspecular reflection, or scattering, occurs when the incident beam strikes boundaries that have irregular surface or are smaller than the beam’s dimensions, resulting in the beam’s energy scattering in multiple different directions (see Figure 1-2). The beam travels around even smaller obstacles without scattering (diffraction). Because a higher frequency results in smaller beam dimensions, obstacles diffracting at lower frequencies act as scatterers at higher frequencies. This explains both higher imaging resolution and higher beam attenuation at higher frequencies. Refraction is the redirection of a beam when striking obliquely at a boundary between two media with different propagation speeds. Unlike reflection, refraction does not contribute to the image formation process but contributes to attenuation (see Figure 1-2). Part of the ultrasound beam’s energy is transferred to the medium in the form of heat. This is absorption, which also increases proportionally to frequency in soft tissues. The bones absorb ultrasound more intensely, together with other energy loss mechanisms, producing acoustic shadows behind them. Finally, part of the original beam is converted by tissues to waves with double or higher-order frequency (harmonic waves). The total propagation losses from the combined effects of scattering, refraction, and absorption are called attenuation, which is directly proportional to frequency. Body compartments with low attenuation that allow imaging deeper structures through them are good acoustic windows (e.g., liquid cavities), while those with high attenuation are acoustic barriers (e.g., bones). The near-total loss of ultrasound at boundaries between tissues and gas makes gas the strongest barrier; nevertheless, important lung ultrasound techniques rely on the abundant artifacts that the aerated lung creates.

Figure 1-1 Propagation speed is different in different tissues (top); electrical impulses stimulate the lead zirconate titanate (PZT) crystal to produce a beam (bottom), while every time an echo is reflected back, the crystal deforms and vibrates, generating another impulse that is processed into an image.

Figure 1-2 Left panel, Reflection (top), specular reflection (middle), and refraction (bottom) of the incident beam. Right panel, Scattering occurs when the incident beam strikes boundaries that are irregular in shape (top) or smaller than the beam’s dimensions (bottom), resulting in the beam’s energy scattering in multiple different directions.

Equipment and imaging modes

Equipment

Ultrasound machines consist of electric pulse generators, transducers, systems for processing received echoes, and image display screens. Modern systems use digital technology and have central processing units running advanced software that forms beams and processes echoes and thereafter stores images. The key elements of transducers (probes) are PZT crystals, matching layers, backing material, cases, and electrical cables (Figure 1 E-1). Modern electronic transducers generate a range of frequencies (bandwidth) around the central frequency, and contain multiple crystal elements (arrays). This permits them to display the sequence of two-dimensional (2D) images so rapidly that motion is displayed as it actually occurs (real-time scanning). Main transducer types are phased array (sector), linear array, and curved array (Figure 1-3). Sector (phased array) transducers (2 to 4 MHz) have small footprints that produce images of sector format through small acoustic windows (e.g., cardiac and cranial applications). Linear array transducers (7 to 15 MHz) provide images in rectangular or trapezoidal format. They feature high resolution and shallow depth of view because their penetration into deeper structures is limited. Convex (curved array, curvilinear) transducers (2 to 6 MHz) of different shapes and sizes produce images in a sector-shaped format with a wide apex. Microconvex transducers (3 to 8 MHz) feature small footprints and are useful in difficult-access areas, such as the neonatal brain. Transducers generating frequencies of 2.5 to 5 MHz feature a larger curvature radius and are used for abdominal imaging. A variety of convex arrays operating at higher frequencies are used in intracavitary and transesophageal scanning. Finally, transducers with frequencies up to 50 MHz are used for endovascular applications and ultrasound biomicroscopy.^1–7

Figure 1-3 Main types of transducers and formats of their produced images (top and middle); basic imaging modes (bottom).

Figure 1 E-1 The basic components of a transducer.

Notwithstanding the similarities of all general-purpose ultrasound systems, it is critical for every user to be especially familiar with a specific machine’s features, transducer choices, and controls in advance and to practice with it sufficiently in nonemergency settings. Attempting to navigate screens or modes while resuscitating a critically ill patient can be a frustrating process. A demonstration of a common ultrasound machine and its essential controls is provided in Video 1-1. The ability to switch transducers between imaging applications or their components helps optimize image acquisition; even seemingly simple examinations, such as the extended Focused Assessment by Sonography for Trauma (e-FAST) evaluation, may require a transducer change, as well as the use of depth and gain adjustment and image optimization techniques. An e-FAST examination using several transducers is demonstrated in Video 1-2, whereas Video 1-3 shows the use of various imaging modes and may be helpful for novice ultrasound users. Some imaging modes may appear to be the prerogative of advanced users; however, novices quickly learn taking advantage of the additional information they offer. Of note, correct choice of transducers and machine settings will help ensure proper identification of pathology or estimation of physiologic parameters, whereas poor preparation may render the study ambiguous or completely nondiagnostic.

Imaging modes (see figure 1-3)

A-mode (amplitude) is a nonimaging mode no longer used in general-purpose machines. B-mode (brightness) is the main imaging mode of any ultrasound machine. Each grayscale tomographic image in B-mode is composed of pixels with brightness, that depends on the intensity of the echo received from the corresponding location in the body. M-mode (motion) displays the movement of structures along a single line (axis of the ultrasound beam) chosen by the operator (Figure 1-4). M-mode is used in the intensive care unit (ICU) for evaluating heart wall or valve motion (echocardiography), hemodynamic status (vena cava analysis), and documentation of lung sliding or movement of the diaphragm. Doppler modes detect frequency shifts created by sound reflections off a moving target (Doppler effect). A moving reflector or scatterer changes the frequency of the beam (Doppler shift), as in (F_s − F_t) = 2VF_t cos Φ/c, where V = the velocity of moving blood cells, c = the propagation speed, F_t = the frequency emitted by the transducer, F_s = backscattered frequency returning to the transducer, and Φ = the angle between beam and blood flow direction). If the beam lines up in parallel with blood flow (Φ = 0 degrees), cos 0 degrees = 1 (maximum Doppler shift). If the beam is perpendicular to the blood flow (Φ = 90 degrees), velocity measurements cannot be performed because cos 90 degrees = 0 (no Doppler shift). Angles in the 45- to 60-degree range are generally preferred.⁷