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Chapter 13: THE ECHOCARDIOGRAM PRINCIPLES OF ECHOCARDIOGRAPHY Physics and Instrumentation

Sound is an energy form that travels through a medium as a series of alternating compressions and rarefactions of the molecules (Fig. 13-1). Sound is typically characterized by its wavelength, which is the distance between any two consecutive phases of the cycle (e.g., peak compression to peak compression), and by its frequency, which is the number of wavelengths per unit time [customarily expressed as cycles per second, or hertz (Hz)]. The velocity of sound is the product of wavelength and frequency; thus there is an inverse relationship between these two characteristics: the greater the frequency, the shorter the wavelength. Ultrasound is sonic energy with a frequency above the audible range of the human ear (greater than 20,000 Hz) and is useful for diagnostic imaging, since, like light, it can be directed as a beam that will obey the laws of reflection and refraction.14-18 Thus, an ultrasound beam will travel in a straight line through a homogeneous medium. If the beam meets an interface of different acoustic impedance, however, part of the energy will be reflected, and the remaining attenuated signal will be transmitted. The reflected energy, or echo, is used to construct an image-in the case of echocardiography, an image of the heart (Fig. 13-2).

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Cycle length

Sine Wave High Resol

Time

Figure 13-1: Sound energy results in alternating compression and rarefaction of particles in a http://cardiology.accessmedicine.com/server-java/Arknoid/amed/hurst/co_chapters/ch013/ch013_p02.html (1 / 12) [2003-1-4 12:37:

conducting 23]

medium. This alternation, which can be plotted against time (or distance), conforms to a sine-wave pattern (bottom panel). (Modified from Hagan AD, DeMaria AN. Clinical Applications of Two-Dimensional Echocardiography and Cardiac Doppler. Boston: Little, Brown; 1989, with permission.)

Compressional Wave Diagram

Figure 13-2: Upper panel: Attenuation of an ultrasound beam emitted from a transducer. There is reflection and progressive loss of energy at each interface encountered. Lower panel: the reflected wavefronts are recorded as signals of varying amplitudes (A mode) via the piezoelectric crystal. (Upper panel modified from Hagan AD, DeMaria AN. Clinical Applications ofTwo-Dimensional Echocardiography and Cardiac Doppler. Boston: Little, Brown; 1989, with permission.)

Figure 13-2: Upper panel: Attenuation of an ultrasound beam emitted from a transducer. There is reflection and progressive loss of energy at each interface encountered. Lower panel: the reflected wavefronts are recorded as signals of varying amplitudes (A mode) via the piezoelectric crystal. (Upper panel modified from Hagan AD, DeMaria AN. Clinical Applications ofTwo-Dimensional Echocardiography and Cardiac Doppler. Boston: Little, Brown; 1989, with permission.)

The most fundamental component of any echocardiography instrument is the transducer, which is responsible for both transmitting and receiving the ultrasound signal. The transducer consists of electrodes and a piezoelectric crystal whose ionic structure results in deformation of shape when exposed to an electric current.18 Thus, piezoelectric crystals are composed of synthetic materials, such as barium titanate, that, when exposed to electric current from the electrodes, alternately expand and contract to create sound waves. When subjected to the mechanical energy of sound returning from a reflecting surface, the same piezoelectric element changes shape, thereby generating an electrical signal detected by the electrodes (Fig. 13-3). Thus the transducer both produces and receives ultrasonic signals.

Figure 13-3: A through D: the basic principle of ultrasonic imaging. The piezoelectric crystal is activated, producing a transmitted pulse (T), which reflects off the interface. The reflected pulse (R) excites the crystal, producing an electric current. As the velocity of the pulse is constant, distance can be calculated based on the transit time. (Because the pulse must travel back and forth from the interface, the time is divided by 2.) (Modified from Weyman AE. Principles and Practice of Echocardiography, 2d ed. Philadelphia: Lea & Febiger; 1994, with permission.)

In the past, echographs have both transmitted and received signals of the same frequency. Recently, harmonic imaging has been implemented, in which ultrasound energy is transmitted at one frequency (fundamental) and received at a higher harmonic of that frequency (usually the first). Tissue harmonic signals are created by alteration of the frequency of the wave as it propagates through the structure.1^ Contrast microbubbles produce harmonics by virtue of resonating (expanding/constricting) in the ultrasound field.186 The net effect of harmonic imaging is to reduce the signal intensity of background noise and enhance that from true tissue (or microbubble) structure, although some blooming of the signals from valves may be observed.1^

As an imaging modality, ultrasound carries with it several unique technical difficulties. Sound energy is poorly transmitted through air and bone, and the ability to record adequate images is dependent upon a thoracic window that gives the interrogating beam adequate access to cardiac structures. The degree to which ultrasonic energy will be reflected when striking an interface of differential impedance is dependent upon how perpendicular the interrogating beam is to the interface. When the ultrasound beam is directed parallel or near parallel to the interface, little or no sound energy will be reflected to the transducer. Therefore poor signal transmission, a nonorthogonal orientation of the ultrasound beam to the surface, and energy attenuation can result in failure to record signals from cardiac structures-a phenomenon referred to as echo dropout.19 Conversely, some structures may be such strong ultrasonic reflectors-being perpendicular to the beam or extremely dense-that sufficient energy returns to the transducer to be reflected and again transmitted into the field. This phenomenon can lead to reverberations, or the reproduction of the echoes of anatomic structures at multiple locations within the image.20 In addition, background noise artifacts, or signals generated from the system rather than tissue, can also be encountered. Finally, since the ultrasound beam diverges with distance from the transducer and always has a finite width, targets lying on the periphery of the beam may be recorded and displayed as if they were located along the central scan line (Fig. 13-4). This problem may be accentuated in the setting of very strong reflectors that result in the formation of side lobes.21 In either case, beam-width problems associated with ultrasound may result in the depiction of targets in erroneous locations and create problems in interpreting the images.22

Near Field And Far Field Ultrasound

Figure 13-4: Upper panel: The transducer emits an ultrasonic beam that has a near field (where the beam is relatively focused) and a far field (where the beam width increases). Lower panel: B-mode diagram showing the effect of beam width. In the near field, the beam reflects off only one of two objects in close proximity to each other. In the far field, however, two similarly positioned objects are both within the beam width. Therefore, lateral resolution is compromised and the objects' positions are misrepresented.

Figure 13-4: Upper panel: The transducer emits an ultrasonic beam that has a near field (where the beam is relatively focused) and a far field (where the beam width increases). Lower panel: B-mode diagram showing the effect of beam width. In the near field, the beam reflects off only one of two objects in close proximity to each other. In the far field, however, two similarly positioned objects are both within the beam width. Therefore, lateral resolution is compromised and the objects' positions are misrepresented.

The construction of a cardiac image from ultrasound signals is based upon computation of the distance between an anatomic structure and the transducer (Fig. 13-3). Thus, an ultrasound beam is produced by a hand-held transducer positioned on the thorax and directed into the heart. This beam will travel in a straight line until it reaches an interface between structures of different acoustic impedance, such as blood and myocardium. At this point, some ultrasonic energy will be reflected (depending on the density of interface), some will be scattered, and some will continue forward. The amplitude of the propagating signal will be attenuated because of the reduction in energy at the interface (Fig. 13-2). The reflected sound waves return to the transducer and form the basis of the echogram. Electronic circuitry within the echograph measures the time interval required for the transit of the ultrasound beam from the transducer to the interface and back again. Since the velocity of sound in soft tissue is constant (approximately 1540 m/s), the instrument can calculate the total distance traveled to and from the reflecting surface as the product of transit time and velocity of sound. Interface location is derived as one-half of the total transit distance, and a signal is depicted on an oscilloscope or video monitor at that point (Fig. 13-3). The amplitude of ultrasonic energy reflected from each target interface is represented by the brightness of the signal that is displayed.

The one-dimensional ultrasonic B- (or brightness) mode scan line resulting from a single transmitted beam is the cornerstone of echocardiographic imaging. In the most basic form of echocardiography, a single scan line produced by a piezoelectric crystal is passed through the heart (G-hH; Fig. 13-5). At each structural interface, ultrasonic energy is reflected back and displayed at the appropriate distance as a signal, whose amplitude represents the acoustic impedance or density of the material encountered. These signals are subsequently displayed as dots, whose brightness is proportional to the amplitude of reflected ultrasonic energy. The distance from the transducer of these B-mode dots changes as the cardiac structures move during the cardiac cycle. Accordingly, if repetitive B-mode scan lines are produced and swept across the screen over time, the movement of the heart can be obtained as a time-motion (or M-mode) recording,23 providing dynamic rather than merely static cardiac images Fig. 13-5). In clinical use, the piezoelectric crystal within the transducer is activated by alternating electric current to transient at a rate of approximately 1000 pulses per second. This same crystal also receives the returning echo reflections and actually spends the great majority of the time (>90 percent) in the "receive" rather than "transmit" mode. Because the beam is confined to a single location and transmits ultrasound signals at the pulse rate of the transducer, M-mode echocardiography provides very high temporal resolution. Importantly, M mode is an excellent modality for timing cardiac events or recording high-velocity motion.

As ultrasound technology advanced and it became possible to determine accurately the spatial orientation of the interrogating beam, multiple B-mode scan lines from different imaging angles were collected and displayed in proper alignment to create a 2D image. As opposed to B- or M-mode recordings, which are unidimensional (on an anterior-posterior axis), 2D echocardiography provides additional information in either superior-inferior or medial-lateral directions. At the current time, M-mode recordings are derived from the 2D images rather than as a stand-alone signal.

Several characteristics of sound energy are of fundamental importance in determining the quality of the images obtained. High-quality images require optimal resolution-that is, the ability to distinguish two individual objects separated in space. Short wavelengths yield excellent resolution in echo imaging, since the shorter the cycle length, the smaller the object that will reflect the signal and be detected by the echo scanner. Since wavelength is inversely related to frequency, transducers that emit a high-frequency signal (3.5 to 7.0 MHz or greater) yield high-resolution images. High-frequency signals also overcome an important limitation of ultrasonic imaging associated with lateral resolution. Since ultrasonic beams diverge as they propagate away from the transducer, the width of the beam can become sufficiently great to encompass multiple targets and diminish resolution (Fig. 13-4). The degree of beam divergence is less with high-frequency sonic energy than with low-frequency signals. The smaller wavelengths associated with high-frequency signals, however, are subject to greater reflection and scattering (therefore substantially higher attenuation) as the beam propagates through tissue. The resultant attenuation is greater than that with low-frequency signals and leads to decreased sensitivity. Therefore, in clinical practice, echocardiographic examinations are performed utilizing the highest-frequency transducer capable of obtaining signals from all potential targets within the ultrasound field.23

M-Mode Echocardiography

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