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Echocardiography essentials: Physics and instrumentation

Echocardiography essentials: Physics and instrumentation
Literature review current through: Aug 2023.
This topic last updated: Jan 30, 2023.

INTRODUCTION — Echocardiography is the primary noninvasive imaging modality for quantitative and qualitative evaluation of cardiac anatomy, physiology, and function [1,2]. Two-dimensional (2D) echocardiography provides tomographic or "thin-slice" imaging. Comprehensive echocardiographic examination typically involves imaging the heart from multiple "viewing" orientations as well as with multiple imaging techniques (ie, 2D, Doppler). An understanding of the fundamental principles of cardiac ultrasound is necessary for the proper acquisition and interpretation of echocardiographic data.

The basic physics principles that apply to echocardiography and its instrumentation are reviewed here. The normal anatomic views and protocols for echocardiography, as well as its clinical uses, are presented separately. (See "Transthoracic echocardiography: Normal cardiac anatomy and tomographic views".)

ULTRASOUND WAVES — 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). In contrast, the human auditory spectrum comprises frequencies between 20 and 20,000 Hz.

Frequency and wavelength are mathematically related to the velocity of the ultrasound beam within the tissue (approximately 1,540,000 mm/sec for human tissue) as indicated by the following equations:

Velocity of blood  =  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.

INTERACTION OF ULTRASOUND WAVES WITH TISSUES — 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.

Reflection — 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.

Scattering — 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.

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

Attenuation — 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 — 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 (figure 1).

In contrast to echocardiographic imaging, continuous-wave Doppler examinations utilize a pair of dedicated crystals: one for continuous transmission; and one for continuous receiving. (See "Principles of Doppler echocardiography".)

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

"Axial" resolution along the length of the ultrasound beam

"Lateral" resolution perpendicular to the ultrasound beam

"Elevational" resolution across 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.

Second harmonic imaging — An ultrasound wave traveling through tissue becomes distorted, which generates additional sound frequencies that are harmonics of the original or fundamental frequency. The sound wave becomes more distorted and produces more harmonics the further it travels through tissue. The technique of second harmonic imaging uses broadband transducers that receive double the transmitted frequency [3,4]. As a result, only the additional generated harmonics are received and the fundamental signal is filtered out; this acts as a form of depth compensation, effectively boosting the signal from deeper structures.

Second harmonic imaging was initially developed to intensify the ultrasound signal during contrast echocardiography, since microbubble contrast agents emit harmonic frequencies when hit by ultrasound energy [5] (see "Contrast echocardiography: Contrast agents, safety, and imaging technique"). It can also help improve visualization of cardiac structures in echocardiography without the use of exogenous contrast. When compared to fundamental imaging, it reduces variations in ultrasound intensity along endocardial and myocardial surfaces, enhancing these structures [3,4]. This method is of particular benefit for patients in whom optimal echocardiographic images are technically difficult to obtain (image 1) [5]. As an example, one study found that second harmonic imaging enhances endocardial visibility and definition during dobutamine echocardiography [6].

IMAGING MODALITIES — Various echocardiographic modalities are in clinical use and each plays an important role in the evaluation of the heart.

M-mode — Motion or "M"-mode echocardiography is among the earliest forms of cardiac ultrasound and used infrequently today in the routine examination except for better discrimination of pathologies (eg, vegetation). 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 are displayed on the monitor at sweep speeds of 50 to 100 mm/sec. Although originally performed "blind" using dedicated crystals, alignment of the M-mode beam is now performed with 2D imaging guidance (image 2).

Though infrequently used today in routine studies, the very high temporal resolution afforded by M-mode imaging permits the identification of subtle abnormalities such as fluttering of the anterior mitral leaflet due to aortic regurgitation or movement of vegetation.

Two-dimensional (2D) imaging — A 2D image is generated from data obtained electronically using a phased-array transducer (image 3A-D). Since each scan line of data requires a finite period of time for transmission and reception, the time required to complete each 2D image is directly related to the number of scan lines. Thus, there is a trade-off between scan line density and image frame rate. For cardiac applications, a high frame rate (at least 25 frames/second; 40 msec frame rate) is desirable for most situations. The signal received undergoes a complex manipulation to form the final image displayed on the monitor including signal amplification, time-gain compensation, filtering, compression and rectification.

Other modalities — Additional echocardiographic modalities include Doppler echocardiography and three-dimensional echocardiography. (See "Principles of Doppler echocardiography" and "Three-dimensional echocardiography".)

BIOEFFECTS AND SAFETY — There are no known adverse effects from diagnostic ultrasound at clinical imaging frequencies. Ultrasound does have significant bioeffects depending upon the intensity of the exposure. These include thermal effects, which predominate, and cavitation.

Thermal effects — As an ultrasound beam passes through tissue, heat is produced due to the absorption of the mechanical energy generated by the sound wave. The rate of temperature increase is related to the absorption coefficient of the tissue for a given frequency, the density of the tissue, specific heat of the tissue, and the intensity of the ultrasound exposure.

Increases in temperature due to ultrasound are offset by blood flow through the tissue and heat diffusion. More dense tissues, such as bone, heat more rapidly than less dense tissue, such as fat. The rise in temperature for soft tissue has been estimated to be as much as 0.6°C/minute of continuous ultrasound exposure. Since ultrasound exposure and transmission is intermittent, exposure time is far less than study time.

Cavitation — Cavitation is the creation or vibration of small gas-filled bodies by the ultrasound beam; it is more common using higher intensity exposures. Microbubbles resonate, expand, and contract depending upon their dimension in relation to the sound wave. Microbubbles can also be created by ultrasound due to expansion of small cavitation nuclei.

The intensity of ultrasound can be expressed as the power per area during a unit time:

Intensity  =  joules/time  x  time/area  =  joules/area (watt/cm2)

The American Institute of Ultrasound in Medicine has recommended limits for ultrasound exposure. These include:

≤1°C elevation of tissue temperature

≤100 mW/cm2 with unfocused and ≤1 W/cm2 with focused ultrasound beams

A major limitation of measuring the intensity of ultrasound exposure is that transmission is intermittent and the ultrasound beam does not penetrate the same area for prolonged periods.

SUMMARY AND RECOMMENDATIONS — An understanding of the fundamental principles of cardiac ultrasound is necessary for the proper acquisition and interpretation of echocardiographic data.

Sound wave characteristics include frequency, wavelength, and amplitude. The resolution of a recording (the ability to distinguish two objects that are spatially close together) varies directly with the frequency and inversely with the wavelength. Imaging with higher frequency (and lower wavelength) transducers permits enhanced spatial resolution. However, because of attenuation, the trade-off for use of higher frequency transducers is reduced tissue penetration. (See 'Ultrasound waves' above.)

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. (See 'Interaction of ultrasound waves with tissues' above.)

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. (See 'Ultrasound transducers' above.)

Image quality with 2D echocardiography is affected by axial resolution, lateral resolution, and elevational resolution. (See 'Resolution' above.)

There are no known adverse effects from diagnostic ultrasound at clinical imaging frequencies. (See 'Bioeffects and safety' above.)

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