Ultrasound (US) uses the transmission and reflection of mechanical waves to generate an image. US frequency (f) is the number of wavelengths per second and is measured in hertz (Hz) (Fig. 2-1). As humans can hear sound in the 20–20,000 Hz range, US is defined as sound with a frequency >20,000 Hz, or 20 kilohertz (kHz). Diagnostic sonography for most medical applications uses frequencies of 2–20 megahertz (MHz).
US frequency is independent of the medium through which the sound is traveling and is a property of the crystals in the US transducer. Modern transducers now often include a range of frequencies (“broadband”) and/or allow for frequency adjustment. Propagation speed (c) describes how fast the US travels through a given medium. Unlike frequency, the propagation speed depends on the medium through which the sound travels. The wave velocity in fluid or tissue is approximately 1540 m/s (although it does vary slightly depending on the type of tissue) as compared to the propagation speed through air at a velocity of approximately 330 m/s. The relationship of frequency, wavelength, and propagation speed is described by the equation c = f λ. Because frequency is constant, wavelength varies directly with propagation speed (ie, f = c/λ). Thus, when propagation speed increases, wavelength increases, and vice versa. The change in the speed of sound at tissue interfaces results in a change of wavelength, which is responsible for determining image contrast, and resolution.
The power of an US wave refers to the amount of energy passing through the tissue in a unit of time and is expressed in watts. In the majority of compact US units, the power is fixed, though it may be adjusted on more sophisticated machines. If using one of these units, one should always use the lowest power that will produce the desired imaging as higher powers (>1 W) can result in cellular and tissue damage. This principle is commonly referred to as ALARA (As Low As Reasonably Achievable).
Generating the Ultrasound Image
Transducers convert one form of energy into another. Piezoelectric transducers convert electrical energy into mechanical energy by inducing vibration of the ferroelectric materials in the transducer head. These vibrations are transmitted through the tissue, echo back at boundaries of tissue that have different acoustic impedance, and are then converted back to an electrical signal. The transducer thus acts as an US transmitter and receiver. When the boundary between two tissues has high acoustic impedance, most of the US is reflected back to the transducer. If two materials have the same acoustic impedance, their boundary will not produce an echo. Typically, only a small fraction of the US pulse is reflected back, with the majority of the pulse continuing along the beam line, but can also be scattered or refracted.
As an US pulse (or echo) propagates through tissue, the energy contained within the beam progressively diminishes, or becomes attenuated. Attenuation results from absorption of energy, with the energy lost as heat, as well as from scattering of the US beam. The amount of attenuation depends on the frequency, as well as the medium through which the US beam travels. In soft tissue, US energy is absorbed and scattered, and the amount of attenuation is directly proportional to the frequency. Though liquids do not significantly absorb or scatter US energy, attenuation does occur, and is proportional to the square of the frequency. Therefore, to image structures deep in the body, lower frequency transducers are required. Higher frequency waves, however, have better axial resolution, or the ability to distinguish two objects along the beam axis. Lateral resolution depends on the beam width as well as the focal zone (see later) of the transducer. The ideal transducer would be of sufficient frequency to penetrate to the desired depth without significant attenuation, yet have a frequency high enough to provide the best resolution (Fig. 2-2).
Appearance of Ultrasound Images on the Screen
The received echo signal is displayed visually in either a brightness mode (B-mode) (Fig. 2-3, top) or a motion mode (M-mode) (Fig. 2-3, bottom). M-mode displays the motion (Y-axis) of the echo reflection over time (X-axis) on a single line of the B-mode image and is used for precise measurements of size and distances, especially with rapidly moving structures such as cardiac valves or fetal heart rate. In both B- and M-modes, it is the amplitude (measured in decibels) of the returning echo signal that determines the brightness on the screen, and the amount of reflected echo is a function of the density and nature of the target, as well as the angle at which the sound wave is reflected. The B-mode image is produced by sweeping the US pulse perpendicular to the axis of the US beam, and as this occurs at rates of 20–40 frames per second, it is seen as a continuous image by the human eye.