The piezoelectric crystal within the US transducer acts as both a signal transmitter and signal receiver. Most current commercial transducers use ceramics for the piezoelectric material. When an alternating electric current is applied to the piezoelectric crystal, it rapidly vibrates, emitting a high-frequency US as it expands and contracts. The transducer emits a short burst of US, lasting 1–6 μs, called the pulse length. The transducer then enters into a receiver mode referred to as “dead time,” waiting for reflected US waves to return. The combined duration of a single pulse length time and dead time is known as the pulse repetition frequency. When a reflected US wave returns to the crystal, it induces the crystalline generation of an electric current that is detected. This transmit-receive cycle is repeated temporally and spatially, and the differences in time delay of the reflected sound waves received by the transducer are used to generate the US image. To create a 2-dimensional echocardiogram sector image, the transducer must emit and receive at repetition rates of 3,000–5,000 times/s.

The simplest type of US imaging is based on the alternating pulsed transmission and reception along a single US beam. This is used in M-mode imaging, when the image obtained along a single US beam is recorded over time in what is sometimes referred to as the “ice pick” view (Fig. 6.2a). A phased array sector scanner must be used to obtain the familiar 2-dimensional (2D) echo images. In phased array scanners, numerous small piezoelectric crystals are interconnected electrically and arranged in an array. Each crystal is connected to an electrode, which enables each crystal to transmit and receive US signals. The electrical excitation of each crystal within the array is timed such that excitation for every crystal does not occur simultaneously but rather in a timed sequential excitation pattern, allowing the US beams to be “steered” through a sector arc to generate a 2D image (Fig. 6.2b).

During the transmission of US waves, the frequency of the original wave may be altered by nonlinear tissue interactions. These interactions result in the generation of new frequencies that were not present in the original transmitted US beam. These new frequencies are integer multiples of the original signal and are known as harmonic frequencies. If the original sound wave has a frequency of f, the harmonic frequencies are 2f, 3f, etc. The original US beam undergoes significant attenuation as it propagates through the patient’s tissue but generates new harmonic waves with higher-frequency oscillations than the original US beam. These higher-frequency waves have higher spatial resolution and increase in strength as the US wave penetrates the body. Harmonic frequencies are optimal at depths of 4–8 cm which is ideal for imaging cardiac structures. Additionally, strong US waves create strong harmonic frequencies, while weak US waves produce almost no harmonic frequencies. In tissue harmonic imaging, the reflected signal from the original US beam is suppressed, and the harmonic frequencies are used to construct the image. The development of tissue harmonic imaging has been instrumental in significantly improving echo image quality, particularly in the visual assessment of endocardial borders.

Doppler echo is the detection of signals from anatomic structures with a lateral dimension smaller than the wavelength of US. Doppler echo is integral in the echocardiographic measurement of blood velocities, tissue velocities, and in newer techniques measuring ventricular torsion and strains. For a stationary anatomic structure, any US waves that strike the anatomic structure and reflect back to the transducer will have the same frequency as the originally transmitted US beam. If the anatomic structure is in motion in a direction away from the transducer, the frequency of the reflected beams will be lower than the originally transmitted frequency. If the anatomic structure is moving toward the transducer, the frequency of the reflected beams will be faster than the originally transmitted signal frequency. This is the same sound wave phenomenon observed when a train moving toward an observer has a high-pitched whistle, which then changes to a low-pitched whistle as the train travels away from the observer. The difference in frequencies between the originally transmitted frequency and the reflected, returning frequency is known as the Doppler shift. In echo, it is possible to calculate the velocity of blood as it travels through cardiac structures by detecting the Doppler shift from the scattered echo’s reflected from the thousands of blood cells located within the US beam. The velocity of the moving blood cells can be calculated using the equation:V = c(F_s – Fτ)/ (2Fτ ‑cosθ)

F _τ is the transmitted frequency, F _s is the frequency of the signal returning from the blood, c is the velocity of sound waves in blood (1,540 m/s), and cosθ is the intercept angle between the US beam and the direction of blood flow. In cardiac applications, care must be taken to align the transducer parallel to the direction of blood flow so that cosθ = 1. Although small angle deviations typically result in minor underestimations of blood velocity, an angle deviation of 60° between the US beam and the direction of blood flow can result in a 50 % underestimation of blood velocity.

There are three types of Doppler interrogations performed during a typical echocardiogram study: continuous wave (CW) Doppler, pulse wave (PW) Doppler, and color Doppler (Fig. 6.3). In CW Doppler, two crystals are used. One crystal is set to continuously transmit an US signal, and the other crystal is set to continuously receive return signals. Along the single US beam, every Doppler shift is recorded simultaneously and displayed in graph format as velocity versus time. The advantage of this method is that high velocities can be measured accurately due to the continuous US signal transmission and data reception. In contrast to CW Doppler which samples all velocities along the beam length, PW Doppler allows for the measurement of velocity at a prespecified sample depth. A pulse of US frequency is transmitted to the sample site, and after a specific time period, the receiver detects the reflected signal. The specified time interval is determined by the time it takes the US to reach the sample depth and return and is called the pulse repetition frequency. Color flow imaging is conceptually analogous to PW Doppler, but instead of a single sample point, numerous sample points along multiple sampling lines are obtained. By combining data points among the different sample sites, mean blood flow velocities are calculated and a 2D image of blood flow is constructed. By convention, flow toward the transducer is depicted in red, and flow away from the transducer is depicted in blue. A third color, typically yellow or green, may be used to indicate “variance” or excessive variation in velocity within the sample points to represent nonlaminar or high-velocity blood flow.

The incorporation of echo into the evaluation of CHD has changed the course of management for patients born with heart defects [3–5]. Prior to widespread use of 2D echo, diagnosis and follow-up evaluation of CHD was done primarily by history, physical exam, chest X-ray, electrocardiogram, and cardiac catheterization. There is still no replacement for a detailed history and physical exam, but a detailed anatomic and hemodynamic evaluation is necessary upon initial diagnosis to properly plan for both initial intervention and subsequent follow-up care. Cardiac catheterization historically had been the only diagnostic tool to acquire anatomic, functional, and hemodynamic data to guide management. While the information derived from catheterization was incredibly valuable, its invasive nature and lack of portability remained its biggest disadvantages [6, 7]. Technological advances have allowed the incorporation of US imaging as a noninvasive method of evaluating patients with CHD. The high temporal and spatial resolution of echo enables it to provide accurate and comprehensive anatomic and functional diagnostic assessments that, in turn, guide the management and follow-up of patients. For a large proportion of patients, a high-quality echo study provides sufficient information to recommend specific surgical or percutaneous therapies without the need for more invasive diagnostics [8, 9]. Today, it is rare for patients with even the most complex cardiac defects to require catheterization prior to initial surgery.