Since the late 1970s, perhaps the most important change in the practice of clinical paediatric cardiology has been the introduction of cross sectional echocardiography, with this advance then being facilitated by ongoing and continuous technical improvements. Echocardiography currently allows highly accurate diagnosis of nearly all morphological abnormalities, for the most part making invasive diagnostic techniques obsolete. Nowadays, cardiac catheterisation and angiography is no longer a routine investigation in most patients prior to cardiac surgery, as it used to be a few decades ago. Indeed, the main indication for such an investigation, apart from the purposes of intervention, is the display of structures beyond echocardiographic visualisation, such as peripheral pulmonary arteries, distal coronary arteries, and more rarely complex abnormalities of systemic and pulmonary venous return. For these indications, magnetic resonance and computerised tomographic imaging are progressively replacing cardiac catheterisation. All paediatric cardiologists, cardiac intensivists, and paediatric cardiac surgeons, therefore, should be familiar with echocardiographic imaging and diagnosis. The first part of this chapter is devoted to the principles of cross sectional and Doppler echocardiography as applied in clinical practice. In the subsequent parts of the chapter, our colleagues discuss some more advanced applications, such as tissue Doppler techniques, three-dimensional echocardiography, and computed tomography and magnetic resonance imaging.
PHYSICAL PRINCIPLES OF ULTRASONIC IMAGING
We begin with a limited introduction to the physical principles of ultrasound, as some background knowledge is relevant for optimization of images in clinical practice. Those interested in more details on the physics of imaging should consult more comprehensive reviews, or specialised textbooks of echocardiography. 1
The Physical Properties of Ultrasound
An acoustic wave is a mechanical wave causing particles to be displaced from a position of equilibrium while traveling through the medium, causing local compression and rarefaction influenced by the elasticity and density of the medium. Ultrasonic waves are longitudinal sound waves with a frequency above 20 kHz, the highest frequency which can be detected by the human ear. The waves are generated and detected by a piezoelectric crystal, which deforms under the influence of an electrical field. The velocity of a sound wave is dependent upon the density and stiffness of the medium. Stiffness is the hardness or resistance of the material to compression. Density is the concentration of the matter. Increase in stiffness also increases speed, whereas an increase in density decreases speed. Transmission of such sonic waves is slow in air or gasses, but fast in solid mediums. In the tissues of the body, the velocity of sound is relatively constant, at 1540 m/sec. The frequency represents the numbers of cycles occurring in each second of time, and is expressed in hertz, with 1 hertz corresponding to one cycle per second. The wavelength corresponds to the length over space over which one cycle occurs. The amplitude reflects the strength divided by the intensity of a soundwave, and is expressed in decibels. The velocity of transmission, the frequency, and the wavelength are related by the formula c = f × λ, where c is the speed of sound through the medium, f is the the frequency of the wave, and λ is the wavelength.
The range of frequencies used for medical applications is from 2 to 12 megahertz (mHz). Corresponding wavelengths, therefore, are in the range 0.8 to 0.13 mm. This implies an intrinsic limitation for spatial resolution, as two structures need to be separated by more than one wavelength in order to be resolved. This relationship also explains why a probe with a higher frequency has a greater spatial resolution. When sound travels through a homogeneous medium, different interactions occur, such as reflection, attenuation, refraction and scattering. Reflection occurs at a boundary or interface between two mediums having a different acoustic density. The difference in acoustic impedance, known as z , between the two tissues causes reflection of the sound wave in the direction of the transducer. Reflection from a smooth or specular interface between tissues causes the sound wave to return to the transducer. Irregular interfaces will cause scatter in different directions. The ultrasonic image is created based on the reflected waves. Refraction is the phenomenon that the ultrasound wave, which passes through the tissue, is refracted based on the incidence of the beam. The incidence is oblique when the direction of the sound beam is not at a right angle to the boundary of the two mediums. This same phenomenon explains why a straight pencil that sits in a glass of water appears to have a bend in it. Attenuation is the loss of sonic energy as sound propagates through a medium. It is produced by the absorption of the ultrasonic energy by conversion to heat, as well as by reflection and scattering. The deeper the wave travels in the body, the weaker it becomes. The amplitude and strength of the wave decreases with increasing depth. Overall attenuation is dependent upon the frequency, such that ultrasound with lower frequencies penetrates more deeply into the body than ultrasound with higher frequencies. Probes with lower frequencies produce better penetration, but as mentioned above, have lower spatial resolution. Probes with higher frequencies have lower penetration, but higher spatial resolution. Attenuation also depends on acoustic impedance, and on mismatch in impedance between adjacent structures. Since air has a very high acoustic impedance, any air between the transducer and the cardiac structures of interest results in substantial attenuation of the signal produced. This is avoided on transthoracic examinations by use of water-soluble gel to form an airless contact between the transducer and the skin. The air-filled lungs are avoided by careful positioning of the patient, and by the use of acoustic windows that allow access of the ultrasonic beam to the cardiac structures without the need to pass through intervening lung tissue. Attenuation causes decrease in amplitude as the wave passes through the tissue. In most ultrasonic systems, this is corrected for by automatic compensation. Further manual compensation in terms of gain in depth or time can be achieved by changing controls on the machine, which correct for the automated attenuation at specific depths of image.
Production of Images
Each ultrasonic pulse, encountering numerous interfaces, gives rise to a series of reflected echoes returning at time intervals corresponding to their depths. In this way, each pulse from the ultrasonic crystal demonstrates a line of information that corresponds to the structures encountered by the sound beam. The returned signal is processed so that the radiofrequency signal is converted into an image. The M-mode represents one line of information displayed on the vertical axis, with time on the horizontal axis. Its advantage is the very high temporal resolution. The frequency of repetition of the pulse producing a typical M-mode trace is approximately 1000 frames per second. The shade of grey on the M-mode recording is determined by the intensity of each reflected echo.
Conventional cross sectional echocardiography depends on the construction of an image using multiple individual lines of information. Typically 64 or 128 lines of information are required to produce one image or frame. Multiple frames are constructed in real time each second, the limiting factor being the time necessary for the echoes from each pulse to return to the transducer. At depths of 5 to 15 centimetres, it is possible to achieve frame rates of 28 to 50 per second. Modern systems are capable of manipulating the frame rates by sending out different pulses at different times, but in general increasing the frame rate will reduce the quality of the image, and hence its spatial resolution.
Quality of Images
When concerning the image, quality refers to the resolution of the imaging system. Spatial resolution refers to the capacity of the system to resolve small structures. It can be considered as the smallest distance by which the system is capable of identifying two dots as separate entities. Contrast resolution is the ability of the system to distinguish differences in the density of the soft tissues. Temporal resolution refers to the capacity of the system to resolve differences in time. There are many potential sources for artifacts in ultrasonic images that affect their quality.
Spatial Resolution
This can be defined as the combination of axial and lateral resolution. Axial, or depth, resolution is the capacity of the ultrasonic system to distinguish how close together two objects can be along the axis of the beam, yet still be distinguished as two separate objects. Wavelength affects axial resolution, and it is improved by increasing the frequency. Axial resolution is much higher compared to lateral resolution, which is the capacity of the system to resolve two adjacent objects that are perpendicular to the axis of the beam as separate entities. The width of the beam affects the lateral resolution: the wider the beam the lower the lateral resolution. This is influenced by the focal zone, which is the depth of the smallest beam width. The near field is the zone between the transducer and the focal zone, and the far field is the region beyond the focal zone. Optimising the focus at a certain depth optimizes lateral resolution. The limit using the focal zone is determined by the size and frequency of the transducer. Small transducers focus well in the near field, while large transducers perform better in the far field. The width of the beam is also influenced by the frequency of the transducer, with higher frequency probes having a better lateral resolution compared to those which have a low frequency. Probes with higher frequency, however, suffer from their limited ability to penetrate into the tissue. Line density can also be improved by decreasing the sector width, resulting in better lateral resolution, but more limited image sector.
Temporal Resolution
To image moving objects, structures such as blood and heart, the frame rate is important, and is related to the motion speed of the object. The eye generally can only see 25 frames per second, providing a temporal resolution of about 40 msec. The temporal resolution is limited by the sweep speed of the echo beam, which in turn is limited by the speed of sound, as the echo from the deepest part of the image has to return before the next pulse is sent out at a different angle in the neighboring beam. The speed of the sweep can be increased by reducing the number of beams, or increasing the beam width in the sector, by using the frame rate control, or by decreasing the width of the sector. The first option decreases the lateral resolution, while the second decreases the image field. Temporal resolution, therefore, cannot be increased without a trade-off, due to the physical limitations of echocardiography.
Ultrasonic Imaging Artifacts
There are many potential sources for artifacts in echocardiography. For interpretation of the images it is important to know and recognise them.
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Drop-out of parallel structures. When structures are parallel to the ultrasonic beam, there is very little reflection caused by the structure, resulting in drop-out. A typical example is the atrial septum when viewed from the apex.
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Acoustic shadowing. Transmission of the ultrasonic beam through the tissue is influenced by the presence of tissue with very high density. Typical examples are prosthetic valves, devices, catheters, or calcifications. These structures can make it impossible to view structures behind them.
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Reverberations can occur with lateral spread of high-intensity echos. Bright echos can have considerable width. A lot of reverberations originate from the interaction between the transducer and the ribs.
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Mirror-imaging. This artifact appears as a display of two images, one real and one artifactual, and is due to the sound beam interacting with a strong reflector.
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Ring down or comet tail. The ring down artifact comes from gas bubble in a fluid medium. The comet tail originates from highly reflective structures, such as surgical clips.
Doppler Echocardiography
The Doppler principle is the frequency shift caused by ultrasonic waves if they are reflected by a moving reflector. The frequency of reflected waves increases as the target approaches the receiver, and decreases as it moves away from the observer. A typical example is the sound of an ambulance siren, which has a higher pitch sound when moving towards the observer, and a lower pitch sound when moving away. When initially applied to echocardiography, the Doppler technique was used to measure the velocities of blood pools. A Doppler shift is caused by interaction of the ultrasonic beam with a moving pool of red blood cells.The Doppler shift depends upon the velocity and direction of the flowing blood, the angle between the beam and the flow, and the velocity of sound within in the tissues. This is expressed in the Doppler formula:
F d = 2 F 0 V cos θ c
where F d is the observed shift in frequency, F 0 is the transmitted frequency, V is the velocity of flow of blood, θ is the angle of intercept between the beam and the direction of blood flow, and c is the velocity of sound in human tissue, which is 1540 m/sec.
The observed frequency shift is several kilohertz in magnitude, and produces an audible signal that can be electronically processed and displayed graphically. As different velocities are present within the ultrasonic beam for any given time instance during the cardiac cycle, a range of Doppler frequencies is typically detected. Thus, a spectrum of Doppler shifts are measured and displayed in the spectrogram, hence the term spectral Doppler as often encountered in the literature. The Doppler-shifted frequencies usually undergo fast-Fourier transformation, converting the original Doppler waveform into a spectral display with velocity on the vertical axis, time on the horizontal axis, and amplitude as shades of grey. Conventionally, Doppler signals from blood moving towards the transducer are displayed above the baseline. Similarly, when blood flows away from the transducer, the display is below the baseline. When the Doppler shift is known, the Doppler equation outlined above can be rearranged to
V = c F d 2 F 0 cos θ
In this way, the velocity of flow can be calculated. The main source of error is in the determination of the angle of intercept between the ultrasonic beam and the axis of flow. When the angle exceeds 20 degrees, it should be measured and included in the Doppler equation. In practice, it is difficult to measure accurately this angle. Potential problems arising from a large angle of interception, therefore, are overcome by obtaining the clearest audio signal with the highest velocity. The ultrasonic beam can then be assumed to be nearly parallel to the direction of flow. As blood velocities cause a Doppler shift in the audible range, the Doppler shift itself can be made audible to the user. A high pitch, with a large Doppler shift, corresponds to a high velocity, whereas a low pitch, with a small shift, corresponds to a low velocity.
Types of Doppler Ultrasound
There are three types of Doppler ultrasound, namely continuous wave, pulsed wave, and colour flow Doppler.
Continuous Wave Doppler
In continuous wave Doppler, one piezoelectric crystal is used for transmitting a continuous wave at a fixed frequency, and a second crystal is used continuously to record the reflected signals. Both crystals are embedded within the same transducer with a slight angle towards each other, and the Doppler shift is continuously sampled. As an ultrasonic wave is transmitted continuously, no spatial information is obtained. Indeed, all velocities occurring anywhere within the ultrasound beam, in other words on the selected ultrasound line of interrogation, will contribute to the reflected signal, and will appear in the spectrogram. As the ultrasound signal weakens with depth due to attenuation, velocities close to the transducer will intrinsically contribute more than the ones occurring further away. The major advantage of continuous wave Doppler is its ability accurately to measure high frequency shifts with no upper limit. It can be used, therefore, to measure high-velocity jets. To optimize the direction of the echobeam within the direction of the flow, cross sectional imaging and color Doppler can be combined with continuous wave Doppler from the same transducer, producing so-called Duplex scanning. This gives a visual impression of being simultaneous, but is achieved by rapid and automatic switching between Doppler and cross sectional imaging. Adding colour Doppler imaging helps optimally to align the beam of interrogation with the direction of the flow of blood.
Pulsed Wave Doppler
Pulsed wave Doppler has been developed to give spatial information on the detected velocities. The technique is not based on the Doppler principle, but provides an output in a spectral display which looks very similar to the way the continuous wave signal is represented. The Doppler shift itself, however, is not measured by the system. When using the pulsed wave system, an image line is chosen along which ultrasonic pulses are transmitted at a constant rate. This rate is the pulse repetition frequency. Instead of continuously sampling the backscattered waves, only one sample of the reflected wave is taken at a fixed time after transmitting a certain pulse. This time interval is the range gate. The range gate will determine the exact depth where velocities are measured. The transducer sends pulses, and must receive that signal back before other pulses can be transmitted. A sample volume is positioned at the area of interest. The velocity of sound in soft tissue is a given constant of 1540 m/sec, and the go and return time is used to determine the depth of the sample volume.
The velocities that can be measured are limited by the pulse repetition frequency, or the number of pulses that are emitted per second. Aliasing of the Doppler signal occurs when the pulse repetition frequency is too low, and the returning signals from one waveform are not received before the next waveform is sent. The frequency at which aliasing occurs is also called the Nyquist limit. Aliasing will result in velocities being displayed at the same time below and above the baseline. The Nyquist limit will be lower the deeper the velocity is measured, as the time for the wave to travel will be longer. The Nyquist limit depends on the frequency of the probe, with higher frequency probes having lower Nyquist limits, and lower frequency probes having higher Nyquist limits. A 2.5 mHz transducer will display velocities of twice the magnitude of those produced by a 5.0 mHz transducer. Appropriate selection of the probe, therefore, is important when performing pulsed Doppler measurements. Different options are possible further to increase the Nyquist limits. The first is to shift the baseline so that velocities are measured in only one direction. A second method is to send out a new pulse before the previous one has returned. This is called the high-repetition frequency method, and results in measuring velocities in more than one site or sample volume, thus reducing the spatial resolution but increasing the Nyquist limit. The size of the sample volume can also be adjusted. A smaller sample volume will result in a sharper velocity profile, as fewer velocities are sampled at the same time.
Colour Flow Interrogation
Colour Doppler displays the direction and flow velocities of the blood superimposed on the cross sectional image. In this technology, different pulses are sent across an image line, and the phase shift between the different signals is measured at two sampling points. This phase shift is proportional to the velocity of the reflecting object. A colour is assigned to the direction of flow according to whether it is away from or towards the transducer. In this respect, it may be helpful to remember the mnemonic BART, indicating blue away and red towards.
The velocity of flow is displayed in shades of these colors. The brighter the colour, the higher is the velocity. The colour information represents the mean velocity of flow. When the flow is disturbed, or not laminar, the pattern will show as a mosaic. This mosaic pattern is only produced if the variance is on. The variance is a colour Doppler option that is present with all the cardiac presets. It is important to appreciate that, because colour flow mapping is a form of pulsed Doppler, it is subject to the same physical principles. Increasing velocities are represented as increasingly bright forms of red or blue until they reach the Nyquist limit, when aliasing then superimposes new colours on the display. Typical Nyquist velocity limits are from 0.6 to 1 m/sec. For flows at high velocity, therefore, multiple aliasing occurs. The colour image is then useful only from a qualitative point of view, and does not permit the demonstration of directional flow. Similarly to pulsed Doppler, the Nyquist limit is dependent on transducer frequency and depth.
Colour flow Doppler provides information regarding the Doppler shift from an entire area, unlike pulsed Doppler, which samples from a specific point. More time is required, therefore, for colour Doppler to compute the lines of information onto the screen. Frame rate and line densities are reduced proportional to the time required. Keeping the colour sector small will provide a better frame rate, and produce a flicker-free image. When using colour flow, the operator is able to visualize the flow of blood in relation to the surrounding structures, which provides a method for rapid interpretation of abnormal location and direction of flow, and helps to guide Doppler interrogation of abnormal flow.
Comparison of Doppler Methods
The three types of Doppler interrogation described are complementary, each measuring the velocities of flow in different ways. For the evaluation of high velocities, the method of choice is continuous wave Doppler, since it does not give rise to aliasing. Although it does not permit gating for precise localisation of the target, a steerable cursor line, and a focused beam, allows precise alignment, thus assuring appropriate measurements of flow. Pulsed Doppler, in contrast, enables measurements of flow at a known depth, allowing more precise calculations, but is limited by the maximal measurable velocity. Colour flow Doppler is a qualitative method, but provides spatial information not obtained with other methods. By permitting visualisation of the disturbed jet, it facilitates the alignment of the continuous wave Doppler beam. The visual effect of colour flow Doppler provides a method of rapidly screening abnormal velocities within the heart, thus directing the more quantitative methods.
EQUIPMENT
M-Mode Imaging
M-mode echocardiography is derived from an M line superimposed on a cross sectional image. The M-mode trace itself shows time as the second dimension. Control of the speed of the sweep enables accurate measurements of intervals in the cardiac cycle, and the high-repetition frequency of the technique allows not only excellent temporal resolution of moving structures, but also precise measurements of mural thickness and cavitary size. In this way, the information derived is superior to that obtained from cross sections. M-mode echocardiography is still commonly used for the evaluation of left ventricular function, using short- or long-axis cuts through the left ventricle, and the timing of cardiac events such as left ventricular ejection time, using a long-axis cut through the aortic valve.
Cross Sectional Imaging
Rather than mechanically moving or tilting the ultrasound crystals, as was the case with the mechanical probes from the previous generation, modern devices make use of electronics to steer the beam. For this purpose, an array of piezoelectric crystals is used. The crystals can be organized in a linear or curved fashion, or more recently in a two-dimensional matrix. By introducing delays between the excitation of different crystals in the array, the ultrasonic wave can be directed without moving the transducer. The signal for transmission in a particular direction is the sum of the signals received by the individual elements. These individual contributions can be filtered, scaled, and time-delayed separately before summing to create the image. This process is referred to as beam-forming. In the newest generation of matrix transducers, a large number of crystals are implanted in a two-dimensional matrix, which enables full steering of the beam in all three dimensions, and is capable of generating three-dimensional images. Currently, a high-frequency matrix transducer has been developed by Philips Medical Systems, which allows high resolution three-dimensional imaging in children. The current resolution of the three-dimensional images produced, however, is significantly when lower compared to the spatial and temporal resolution which can be obtained using the current probes available for cross sectional imaging.
When performing cross sectional imaging, the most appropriate probe must be selected to optimise the quality of the images produced. Probes with higher frequencies, as already discussed, have better spatial resolution but lower penetration, while probes with lower frequencies have a lower spatial resolution but higher penetration. In neonates and small infants high-frequency transducers from 7.5 to 12 mHz provide excellent cross sectional resolution. Older children, in contrast, are better studied with transducers of lower frequencies, from 5 to 3.5 mHz. Harmonic imaging was developed further to optimize the quality of the images, and to reduce the signal-to-noise ratio. Due to the phenomenon of distortion during propagation through the tissues, harmonic frequencies combining integer multiples of the transmitted frequency are generated. Transmitting a pulse of 1.7 mHz, for example, will result in the spontaneous generation of harmonic components of 3.4, 5.1, 6.8, 8.5 mHz, and so on. These harmonic components will grow stronger with the distance of propagation. The scanner can be set up in such a way as to receive only the second harmonic component, thus generating a second harmonic image. Such an image typically produces a better signal-to-noise ratio by avoiding the clutter noise due to reverberation artifacts, such as those produced by the ribs. Such harmonic imaging is often used in patients with poor acoustic windows. Its disadavantage is the reduction produced in axial resolution in the direction of the echobeam. In smaller patients, therefore, it rarely improves the quality of the images, but the technique can be useful in larger children with poor acoustic windows.
Blood-Pool Doppler
Optimising the Settings for Continuous Wave Doppler
The continuous wave image is an echocardiographic image that is influenced by all the parameters that affect a normal cross sectional picture. The gain control affects the ratio of the strength of the output signal to the input signal. The gain controls, therefore, should be manipulated to produce a clean uniform profile, without any blooming. The controls should be increased to over-emphasise the image, and then adjusted down. This will prevent any loss of information because of too little gain. The compress control assigns the varying amplitudes a certain shade of grey. If this control is either very low or very high, the quality of the spectral analysis graph will be affected, and this may lead to erroneous interpretation. The reject control eliminates the smaller amplitude signals that are below a certain threshold. This helps to provide a cleaner image, and may make measurements more obvious. The filter is used to reduce the noise that occurs from reflectors produced from walls and other structures that are within the range of the ultrasonic beam.The volume button should be at the appropriate level to hear the frequencies.
Optimising the Settings for Pulsed Doppler
The frequency of the probe will affect the Nyquist limit of pulsed Doppler, with probes of higher frequency having a lower Nyquist limit. The gain control, compress, and filter settings are similar to those described above for continuous wave Doppler. Shift of the baseline allows the entire display to be used to show either forward or reverse flow, a feature which is useful if the flow is only in one direction. The scale should always be optimised, and be set no higher than necessary to display the measured velocities. The size of the gate should be optimised, with an increase in the sample volume increasing the strength of the signal at the expense of a lower spatial resolution. In general, the smallest sample volume providing an adequate ratio of signal-to-noise should be used. The update function allows for simultaneous duplex imaging to optimise the location of sampling relative to the cross sectional image. Simultaneous cross sectional imaging, however, reduces the temporal resolution. All pulsed-Doppler traces, therefore, should be obtained with the cross sectional image frozen.
Optimising the Settings for Colour Doppler
This involves optimising the gain settings, the scale, and the size of the sector. The gain should be adjusted until background noise is detected in the colour image, and then reducing it so that the noise just disappears. The scale should be adapted depending on the velocities of the flows measured. When looking at flows with high velocities, the scale should be adapted so the maximal Nyquist limit is chosen. When flows of low velocity are studied, such as venous flows or those in the coronary arteries, the scale needs to be lowered. The size of the colour sector needs to be adjusted to optimise frame rates. The smallest necessary size should be used.
STORAGE AND REPORTING OF THE IMAGES CREATED
All current echocardiographic machines allow the recording and storing of images in a digital format, which can be retrieved, viewed, and further analysed on a digital viewing system and work station. This allows easy storage and retrieval of echocardiographic images and studies. More and more centres use digital systems for reporting, which can be integrated within the reviewing stations. The full digital workflow contributes to the quality and efficiency of paediatric echocardiographic laboratories. 2,3
NORMAL CARDIAC ANATOMY
With cross sectional echocardiography, the clinician has at his or her disposal the technique to illustrate cardiac anatomy in all the detail seen by the morphologist, as described in Chapter 2 . To take advantage of the material displayed, a thorough knowledge of normal and abnormal cardiac morphology is essential. It has been rare in the past, however, to find cardiac structures displayed by morphologists as they are oriented within the body. For example, the mitral valve is usually illustrated by incising the atrioventricular junction, and spreading the valvar leaflets and papillary muscles. This gives a spurious impression of the disposition of the papillary muscles. In life, they are adjacent. The echocardiographic sections display the anatomy in cross section. It is helpful, therefore, if the morphologist displays the anatomy in similar fashion. Understanding is helped further by illustrating the anatomical sections in the normal orientation of the heart as it is positioned in the body. 4 In the normal subject, the long axis of the heart is oblique, extending more or less from the left subcostal region to the right shoulder ( Fig. 18A-1 ). The heart does not stand on its apex in St Valentine’s fashion. In addition to the long axis being oblique, the cardiac chambers, particularly the ventricles, are arranged so that the morphologically right structures are anterior to their morphologically left counterparts (see Chapter 2 ). The left atrium is the most posterior of the cardiac chambers. Only its appendage projects to the border of the cardiac silhouette. The right ventricle lies anterior to the left ventricle. It swings across the front of the ventricular mass, reaching from the inferior and right-sided tricuspid valve to the antero-superiorly and leftwardly positioned pulmonary valve. The aortic valve is centrally located within the heart, and is related to all four chambers ( Fig. 18A-2 ). This central position of the aortic valve, located between the tricuspid and mitral valves, is the key to the understanding of cardiac anatomy. Because of the wedged position of the aortic valve, the subaortic outflow tract lifts the leaflets of the mitral valve away from the ventricular septum. Hence, there are no direct attachments of tension apparatus to the muscular septum in the left ventricle. In contrast, the septal leaflet of the tricuspid valve hugs the septum ( Fig. 18A-3 ), and is attached to it by tension apparatus along its length. The short-axis cuts emphasise other significant differences between the right and left sides of the normal heart. The arterial and atrioventricular valves in the right ventricle are separated by the supraventricular crest. No such muscular crest exists in the roof of the left ventricle, where the leaflets of the aortic and mitral valves are in fibrous continuity ( Fig. 18A-4 ). The subaortic outflow tract also has important relationships to the areas of atrioventricular contiguity, these being made up of two components. The first, the atrioventricular membranous septum, is made of fibrous tissue, and is an integral part of the central fibrous body. It interposes between the medial wall of the subaortic outflow tract and the right atrium. The other area is the region of overlapping of the atrial and ventricular musculatures. It exists because the septal leaflet of the tricuspid valve is attached to the septum more towards the ventricular apex than are the leaflets of the mitral valve, producing the characteristic off-setting of the valvar leaflets. In the area between these valvar attachments, the walls of the right atrium overlap the base of the ventricular mass, thus forming a muscular atrioventricular sandwich ( Fig. 18A-5 ). The anatomist describes the location of this area as seen from the diaphragmatic surface of the heart as the cardiac crux, representing the position where the plane of the normal septal structures crosses the inferior atrioventricular groove. The echocardiographer cannot see this point on the epicardial surface of the heart, but is able to identify the so-called echocardiographic crux, the asymmetric cruciate appearance representing the off-set attachment of the atrioventricular valvar leaflets.
Anatomical Principles of Echocardiography
All of the morphological features described above are potentially amenable to dissection with the ultrasonic beam. The echocardiographic tyro will immediately become aware that the ultrasonic beam is obstructed both by bony structures and the air-filled lungs. Because of this, access from the body surface can be limited, although the heart can be viewed from many different aspects. Unobstructed views can be obtained from the cardiac apex, from alongside the sternum through the intercostal spaces, from beneath the rib cage, and from the suprasternal notch (see Fig. 18A-6 ). The cardiac components can also be visualised by placing a transducer within the oesophagus and stomach. From these various windows, the heart and great vessels, according to their position within the chest, can be cut in different planes. Standard echocardiographic views are well established for patients with normally located hearts. Visualisation and understanding of the congenitally malformed heart, however, is facilitated if the examiner develops a technique of orientation and identification based on points of reference that are internal to the heart and great vessels. For example, the sections obtained from the so-called standard approaches will be of little value when a heart is in the right chest, even though it may be structurally normal. The key to a system of universal value is to assess the heart, whatever its position, in terms of its three orthogonal planes. These planes can be described as being coronal, sagittal, and transverse relative to the heart, but the axis of the heart rarely corresponds to the axis of the body (see Fig. 18A-1 ). Of these cardiac orthogonal planes, two are in the long axis of the heart itself, the other being in the cardiac short axis ( Fig. 18A-7 ). Using cross sectional imaging, these index planes cannot all be obtained from a single echocardiographic window. Simple geometric principles dictate that, from any given window, it is possible to obtain only two of these basic planes, although intermediate cuts can be taken towards the third plane.
In the simplest terms, scanning from the apical window provides the two cardiac long-axis planes. The long-axis coronal plane is at right angles to the inlet part of the ventricular septum, and is conventionally termed the four-chamber plane (see Fig. 18A-3 ). This is the plane that shows the anatomy of the atrioventricular contiguities. When taken through the membranous atrioventricular septum, the section also incorporates part of the subaortic outflow tract. Hence, most of these cuts show more than the four basic cardiac chambers. The designation as a four-chamber plane, nonetheless, is a useful one. The long-axis plane at right angles to the outlet part of the septum can, by analogy to the four-chamber plane, be considered as a two-chamber plane. In most cases it also included part of the right ventricular infundibulum, and illustrates more than the two basic chambers. The long-axis plane basically showing two chambers is also one of the standard planes obtained from the parasternal window ( Fig. 18A-8 ), together with the series of short-axis planes ( Fig. 18A-9 ). When scanning from subcostal and suprasternal windows, it is no longer possible to cut a normally positioned heart in its own axes, although this may be possible when the heart is abnormally positioned. Instead, the usually located heart is cut in paracoronal and parasagittal planes relative to the axes of the body. From subcostal position, such cuts produce sections that are similar to, but differ subtly from, the four-chamber and short-axis planes. For example, a paracoronal cut can, on occasion, replicate a heart with absence of the right atrioventricular connection, even when two normal atrioventricular valves are present. These cuts also show to advantage the location of the membranous septum, and how the inferior part of the muscular ventricular septum separates the right ventricular inlet from the outlet of the left ventricle ( Fig. 18A-10 ). These cuts also show the long axis of the aortic arch, although this is probably better seen from the suprasternal window. Many views must be used to obtain a complete impression of cardiac structures. Despite the need for presentation and description of echocardiographic findings in terms of cuts, the key to successful evaluation is to obtain continuous scans from one view to another. In practice, the echocardiographer builds up a three-dimensional whole from a series of two-dimensional parts. The experienced investigator will use any view, no matter how unorthodox, to attain this objective.
Transthoracic Echocardiography
A full echocardiographic study involves a complete description of cardiac anatomy, valvar function, and cardiac systolic and diastolic function. Transthoracic windows are excellent in the majority of infants and children, so that interrogation from these windows can provide all anatomical, haemodynamic, and functional information required for diagnosis and treatment of most of the congenital and acquired cardiac lesions encountered by the paediatric echocardiographer. The majority of children are currently referred for cardiac surgery based only on transthoracic echocardiographic studies. For those aged between 3 months and 3 years, however, lack of co-operation can be a limiting factor, and sedation is generally required to permit performance of an adequate echocardiographic study. Transoesophageal imaging in children is mostly limited to peri-operative imaging.
The description of cardiac anatomy is best accomplished by using the segmental anatomical approach. By combining different echocardiographic windows, it is possible to recognise the precise anatomy of each cardiac segment, along with their interconnections and relations. In this chapter, we concentrate on normal findings. The echocardiographic images of congenitally malformed hearts are described in other chapters. A detailed discussion of functional echocardiography is beyond our present scope, so we will focus mainly on the description of cardiac anatomy, albeit discussing valvar function.
When starting a full cross sectional study in a new patient, most echocardiographers will prefer to start with subcostal imaging, as this allows inferential determination of the arrangement of the organs. The study then proceeds with obtaining parasternal long-axis, short-axis, apical and suprasternal views. A study performed for the purposes of follow-up will often start with parasternal long-axis views, deferring the subcostal imaging until the end of the examination. To help the echocardiographer to reconstruct the three-dimensional anatomy, two-dimensional sweeps can be recorded using the different cuts. For instance, a subcostal coronal sweep from posterior to anterior provides extra information in the third dimension. All images should be presented in their correct spatial, or attitudinally appropriate, position on the screen. The anterior and superior structures, therefore, are displayed at the top of the screen, and the rightward structures are generally placed on the left side of the display, with the exception of the parasternal long-axis cut when, by convention, the cardiac apex is displayed on the left of the screen. The standard views required for a paediatric echocardiographic study, as defined by the American Society of Echocardiography, are all obtained as part of a routine examination. 5 These are the subcostal, apical, parasternal, suprasternal notch and right parasternal cuts.
Subcostal Views
Different subcostal views can be obtained. Typical cuts are obtained in a coronal plane, giving long-axis views, in the sagittal plane producing short-axis cuts, and in the transverse plane. Subcostal imaging begins with a transverse section to determine the arrangement of the abdominal organs by inference from the location of the abdominal aorta and inferior caval vein relative to the spine ( Fig. 18A-11 ). 6 The abdominal vessels can also be imaged in a long-axis view by rotating the probe 90 degrees counterclockwise. Colour imaging will help to distinguish the identity of the abdominal great vessels. Posterior coronal long-axis sections, with the transducer directed to the left of the midline, provide excellent views of the atrial septum ( Fig. 18A-12 ). As the transducer is tilted anteriorly, the superior caval vein, the ventricles, and the left ventricular outflow tract can be imaged ( Fig. 18A-13 ). Further anterior tilting of the probe provides views of the right ventricular ouflow tract. Counter-clockwise rotation of the transducer, with the notch in the transducer positioned inferiorly at 6 o’clock, will produce different cuts as the transducer is swept from right to left in a parasagittal plane. A cut though the atriums allows imaging of the superior and inferior caval veins and the atrial septum ( Fig. 18A-14 ). As the transducer is moved towards the left, the subcostal short-axis section of the left ventricle and the right ventricular outflow tract can be imaged ( Figs. 18A-15 and 16 ). In this view, the right ventricular apex and outflow tract are seen, together with the pulmonary valve and proximal pulmonary trunk (see Fig. 18A-15 ). By tilting the transducer further to the left, the midapical and apical portions of both ventricles can be imaged, along with the corresponding portions of the ventricular septum.
