2.1. Ultrasound Imaging of Tissues
Echocardiography uses the properties of sound waves to differentiate tissues of varied density in the human body. Sound travels in mechanical waves with a speed dependent on the density and elastic properties of the medium in which they are traveling (2). This property of tissue is termed its acoustic density. Ultrasound waves, which are used in medical applications, have frequencies that are higher than those audible to the human ear. ultrasound frequencies are generally over 20,000 cycles per second, or hertz, and most cardiac applications are performed using frequencies of 2 million to 10 million hertz, or 210 megahertz (MHz).
When a sound wave, which is generated by electrical stimulation of a piezoelectric crystal, travels through an interface between two tissues of varied acoustic density, such as myocardium and blood, a portion of the energy is reflected backward (the reflected wave), and the rest travels forward through the next tissue (the refracted wave). The reflected wave is received by the transducer, turned back into electrical energy, amplified, and displayed (1). If there is too much variance between the acoustic density of the tissues imaged (as in air-filled lung and myocardium or bone and myocardium), the entire ultrasound wave is reflected, and the cardiac structures cannot be imaged (1).
The amount of reflected wave detected during ultrasound imaging depends not only on the acoustic characteristics of the interface, but also on the angle of incidence or interrogation. An ultrasound beam that encounters a flat surface perpendicular to the beam will reflect a wave in the direction of the transmitted sound. In contrast, a beam parallel to a structure or that encounters an irregularly shaped structure, as is common in tissue imaging, will be reflected with a degree of scatter that is proportional to the angle of incidence (2).
The ability of cardiac ultrasound to provide anatomical resolution depends on the wavelength of the sound used. The speed of transmission, the frequency, and the wavelength are related by the equation c = f x X or X = c/f where c is the speed of sound in the medium, f is the frequency of the wave (in hertz or cycles per second), and X is the wavelength. Thus, a higher frequency transducer will produce a smaller wavelength and improved resolution along the path of the beam, also termed axial resolution (1).
According to Geva (1), axial resolution is generally two times the wavelength used, so that a 3.5-MHz transducer has a wavelength of 0.43 mm and an axial resolution of 0.86 mm, and a 7.5mHz transducer (commonly used in pediatric imaging) has a wavelength of 0.2 mm and axial resolution of approx 0.4 mm. Unfortunately, the use of high-frequency transducers is limited because the smaller wavelength cannot penetrate as deeply into tissue, and they are therefore less useful for cardiac imaging in adults. Lateral resolution in echocardiography is impacted by the diameter of the beam width, which is a function of the transducer size, shape, and focal plane as well as the frequency (1).
M-mode, or motion-mode, echocardiography was the first type of ultrasound used for clinical cardiovascular imaging. Its use today is primarily limited to assessment of valve motion and reliable reproducible measurements of chamber sizes and function (1,3). In M-mode echocardiography, a narrow ultrasound beam is pulsed rapidly in a single plane through the heart, and the movements of the structures in that single plane are plotted against time with very high temporal and axial resolutions.
M-mode echocardiography can be used to assess cardiac wall thickness, aortic root size, chamber sizes, or ventricular function. In general, left ventricular function is quantitated using M-mode by determining the percentage of fractional shortening of the left ventricle, which is calculated using the following equation:
where SF is the shortening fraction, LVEDD is the left ventricular end-diastolic dimension; and LVESD is the left ventricular end-systolic dimension.
Normal values vary with age and range from 35-45% in infants to 28-44% in adolescents and adults (4,5).
Two-dimensional imaging provides an arc of imaging planes by employing multiple ultrasound beams to provide a cross-sectional view of the heart. Currently, 2D imaging provides the majority of information about cardiac structure and function in routine clinical studies. Two-dimensional imaging requires the presence of multiple beams of ultrasound interrogation in a single transducer, and several types of transducers are available to achieve this.
Transducers available for 2D imaging include mechanical (a sweeping or rotating ultrasound beam), phased array (multiple independently controlled sources), and linear array (a line of crystals simultaneously generating a beam of ultrasound). Today, phased-array transducers are most commonly used because of their: (1) small size, (2) ability to provide simultaneous 2D and M-mode or Doppler imaging, and (3) improved control of focal length for a more uniform image throughout the field of view (6).
In addition to using 2D imaging for viewing anatomical detailing, left ventricular function can be quantitated by this mode using an estimated left ventricular ejection fraction. This method, which has been shown to correlate well with angio-graphic estimates of ventricular function, takes advantage of the conical shape of the left ventricle to estimate end-diastolic and end-systolic ventricular volumes from tracings of 2D images using Simpson's biplane rule (3). The ejection fraction is calculated as follows:
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