Echocardiography in the Dog, Cat and Horse, Porciello, 2009

Prévia do material em texto

Echocardiography in the Dog, Cat and Horse 
Francesco Porciello 
Welcome to the online version of Echocardiography in the Dog, Cat and Horse. 
Echocardiography can give wonderful insights into the cardiovascular function of domestic 
animals, but it can be difficult to understand. ECHOCARDIOGRAPHY IN THE DOG, CAT 
AND HORSE will help the general practitioner learn the benifits and limitations of an 
echocardiographic examination, illustrate image acquisition techniques for the beginning 
ultrasonographer, and provide reference values for experienced practitioners. 
Firstly, I would like to express my gratitude for the digital English language edition of the 
Manual of Echocardiography in the Dog, the Cat and the Horse. The digital edition exceeded 
my expectations in terms of ease of use and image quality. I hope that the manual will be of 
help to VINners in their professional development, specifically in improving their 
echocardiographic interpretative skills. 
I examined the entire text and could not identify any section that did not faithfully interpret 
my original thoughts and published text. The faithful reproduction of the Italian text was 
achieved with the cooperation of Dr. Mark Rishniw, who assisted me in both the preparation 
of the original Italian version (where he contributed a chapter on congenital diseases) and 
also in translating the Italian text into English. Other collaborators who contributed chapters 
to the original Italian manual, and to whom I am equally grateful, include my colleague and 
friend, Dr. Francesco Birettoni, who was instrumental in preparing the figures and images to 
the quality observable in both the published and digital editions. Consequently, I would like 
to include both of these cardiologists as collaborators on the title page of the digital edition. 
Additionally, I would like to credit Dr. Rishniw with the co-authorship of the chapter on 
Congenital Cardiac Diseases, and Dr. Birettoni with the co-authorship of the section on feline 
cardiomyopathies in the Chapter on Acquired Cardiac Diseases. 
Finally, I would like to thank Dr. Paul Pion for believing in me and providing me with the 
opportunity to publish this manual in an international forum, thereby bestowing on me the 
honor of having my work read and used throughout the world. I would also like to thank Dr. 
Eliana Poletto, the publisher of the original Italian text, for extending me the opportunity to 
reach a broader audience via the digital edition. 
Perugia, 19 April 2009 
Francesco Porciello 
Chapter 1 - Introduction 
Echocardiograph 
224 
In 1950, Keidel used ultrasound to examine the heart, and by the mid-50's, together with 
Elder and Hertz, he established the basis for ultrasound techniques to describe certain aspects 
of cardiac anatomy and systolic and diastolic cardiac function. In subsequent years, Holmes 
popularized the use of echocardiography in human medicine in the United States. The 
technique was initially used for the appraisal of mitral stenosis, but its application in the 
diagnosis of pericardial effusion and in the evaluation of the cardiac chambers evoked an 
enthusiastic response from much of the scientific community. Coincidentally, experiments 
with ultrasound provided the first two-dimensional images of the canine and feline abdominal 
organs. 
Echocardiography was introduced into veterinary medicine over two decades ago. It offered a 
non-invasive method of examining the structure and function of the heart and large vessels in 
animals, allowing investigators to obtain images that could be utilized in both clinical 
practice and research. 
This technique, despite its limitations (due mainly to the difficulties of transferring 
methodology standardized in humans to domestic animals), provides an extremely useful 
addition in evaluating cardiac pathophysiology. The ability to observe dynamic images of the 
heart allows the echocardiogram to complement the physical examination and the recording 
of a detailed history. In fact, echocardiography relies on the sequential collection and 
interpretation of the basic diagnostic procedures (such as a thorough physical examination 
and history) to prevent imprecise or misleading diagnoses, and is used in conjunction with 
these other diagnostic steps, rather than as a "short cut " to the diagnosis. 
Of fundamental importance is the ability to recognize a "physiological" finding and 
consequently to distinguish physiological variations from pathology. One must always 
remember that a single echocardiographic projection (view) cannot demonstrate all the 
clinically important findings - it is important to confirm anatomical and functional anomalies 
using alternate views as well as performing multiple measurements of cardiac function. It is 
well recognized in the practice of echocardiography that the initial impression obtained from 
a particular view often suggests pathology that is subsequently refuted by additional 
echocardiographic imaging using different views or modalities (e.g. Doppler or color Doppler 
examination). This "rule" of substantiating findings with a complete echocardiographic 
examination must be adhered to by all diagnostic echocardiographers in order to maintain 
scientific rigor and accurate clinical interpretations. 
The term echocardiogram refers to a collection of images that use ultrasound to examine the 
heart and to record information in the form of "echoes", which are reflected ultrasonic waves. 
The maximum frequency detectable by the human ear approaches 20,000 cycles/sec (20 
kHz). The frequencies used in echocardiography range from 2 to more than 7 million 
cycles/sec (2 to >7 MHz). Techniques originally employed in veterinary echocardiography 
consist of monodimensional (M-mode) and bi-dimensional (B-mode or 2D-mode) imaging, 
which, in recent years, have been augmented by Doppler techniques used for the study of 
blood flow and myocardial kinetics. 
Doppler echocardiography requires equipment that has only recently gained popularity in 
veterinary medicine, due to the fact that equipment cost has dropped substantially, together 
with technological developments in image processing and acquisition and portability 
allowing marketing of portable units that offer great performance with few problems. Until 
recently, few people understood how to properly apply the more complex echocardiographic 
techniques (namely pulsed wave or continuous wave Doppler), but the scene today is 
decidedly different. Reference intervals have been characterized in many situations and for 
many animal species of interest allowing accurate diagnostic and prognostic interpretation. 
The primary clinical development of echocardiographic techniques (both regular and 
Doppler) for dogs and cats occurred in the 1980s. The first cardiovascular applications of 
ultrasound in the horse however, originated in the latter half of the 1970s, but progress in 
equine imaging has been slower than in small animals because of probe design that is 
primarily directed towards advances in imaging human patients, thereby limiting imaging 
depth and penetrating power. 
Use of both M-mode and B-mode imaging allows acquisition of information about the 
morphology and the dimensions of the cardiac chambers and walls, as well as changes to the 
valves, the parietal and valvular endocardium, the pericardium and pericardial space, the 
structures immediately connected to the heart and the roots of the great vessels. B-mode 
imaging, which provides a realistic image of the heart in motion, lends itself to qualitative 
assessments, while precise linear measurements can be acquired with the M-mode imaging, a 
graphical representation of the movements of the various structures imaged over time. 
However, with newer processing technology, many of the limitations with B-mode imaging 
(specifically low frame-rates)have been overcome and many traditional M-mode calculations 
can be adequately obtained from 2D images. 
The development of Doppler techniques through the late 1980s provided additional 
diagnostic capabilities that allow appraisal of blood flow blood within vessels and cardiac 
chambers. Specifically, this methodology evaluates five fundamental characteristics of blood 
flow: direction, quality (laminar or turbulent), speed, timing and location. Several types of 
Doppler echocardiography exist: spectral Doppler (comprised of continuous wave and pulsed 
wave technologies), and color Doppler, both of which (like standard M-mode and B-mode 
imaging) involve the emission and reception of sound waves and various graphical 
representations of these sound waves. However, with Doppler imaging, it is the velocity and 
direction of the sound waves that is of interest, rather than just the location of the reflective 
surface. As with traditional echocardiographic images, Doppler echocardiography allows one 
to obtain "anatomic" (color Doppler) or "graphical" (spectral Doppler) representations of 
blood flow in which the variables of location, timing, quality, speed and direction can be 
determined. With spectral Doppler, these variables are represented within a cartesian system 
of reference. On the other hand, color Doppler echocardiography, introduced to clinical 
veterinary practice in the early 1990s, allows for an immediate survey of blood flow in the 
various regions of the heart and vessels and, compared to spectral Doppler, often provides 
"exciting" images by integrating direction, quality and timing of blood flow with real-time 
two-dimensional images. 
To date, there is no evidence of either acute or chronic adverse biological effects from the use 
of diagnostic ultrasonography on either patients or ultrasonographers, with the exception of 
fetal imaging. Therefore, the benefits of these methodologies in veterinary patients far exceed 
any risks that may be associated with their use. However, it is good practice to minimize the 
time of exposure to ultrasound waves and to use the lowest possible power settings that allow 
diagnostic quality imaging. The more intuitive ultrasound techniques, such as B-mode and 
color Doppler echocardiography, should not be relied upon exclusively, at the omission of 
spectral Doppler or M-mode methods. Rather, the clinician must know the diagnostic 
potential of each imaging technique and be able to choose the most appropriate technique for 
each specific patient, understanding that the information obtained from each 
echocardiographic technique is complementary to the others, and that the most accurate 
diagnosis incorporates the assimilation of information from all these techniques. 
The advantages of echocardiography, compared to other diagnostic imaging techniques, are 
essentially due to its safe and practical nature, and can be inferred from the following salient 
characteristics: 
 The examination is painless for humans and domestic animals alike. Therefore, it can 
usually be performed without sedation and can be repeated frequently because of the absence 
of known acute or cumulative side effects; 
 Imaging is safe, even if practiced on pregnant or young animals; 
 The equipment is often portable, lending itself to examinations "in the field" or outside of a 
specific examination room; 
 A complete M-mode and B-mode examination does not generally require more than 30 
minutes; if a Doppler examination is also performed, the complete study usually takes just 15 
minutes longer; 
 The technique is valid for screening and early diagnosis of subclinical or occult pathologies 
or in monitoring inherited cardiac diseases, allowing clinicians to document changes in the 
clinical status over time; 
 In some specific forms of cardiac disease, the echocardiogram represents the only 
diagnostic test capable of providing diagnostic and prognostic information; 
 Echocardiography is relatively cheap to perform with few overhead costs beyond those 
associated with purchasing or leasing the equipment; it consumes little material with the 
exception of imaging gel and printing paper (and, of course, electricity!) 
Chapter 2 ‐ Acquisition Of The Echocardiographic Image 
SEARCH RESULT #: 1 
TITLE: Principles of Ultrasound Physics 
AUTHOR(S): 
ADDRESS (URL): http://www.vin.com/Members/Proceedings/Proceedings.plx?CID=ECHO2007&PID=20489&O=VIN 
 
Sound waves have been employed for centuries in medicine, for example, in percussion during a physical 
examination, where hands and ears work like a rudimentary echocardiogram -- sounds are sent deep into 
tissues and the returning echoes are analyzed with a stethoscope or naked ear. With echocardiography, or 
ultrasonography, the "resolution" of the percussive examination has been improved by moving the transmitted 
sounds to very high frequency spectra and employing electronic reception and analysis of the reflected sounds 
- basically, "better hands and ears". Specifically, ultrasonography is the graphical representation of the analysis 
of the reflected sound waves, or echoes, generated by transmission and reflection of VHF sound waves through 
tissues. Ultrasonography is based on the elastic property of acoustic waves, which can be stretched and 
compressed, penetrate tissues and reflect from tissue interfaces. These elastic acoustic waves are related to 
the speed of sound by frequency and wavelength according to the formula: 
speed of sound = wavelength x frequency (c = λ x ν) 
A sound wave can only travel through a medium (e.g. air, water, tissue). A single compression and 
expansion of the medium constitutes a single acoustic cycle and multiple cycles constitute an ultrasound wave. 
The term "ultrasound" indicates that the sound waves used for imaging have a frequency that exceeds 20 kHz 
making them imperceptible to the human ear. 
The speed with which these waves propagate through an object is directly proportional to the density of the 
object. In echocardiography, the speed of sound turns out to be almost constant since, in most tissues, the 
speed of sound ranges from 1500 to 1600 m/sec. In clinical ultrasonography, the accepted speed of sound is 
assumed to be 1542 m/sec. The density of objects propagating the sound waves also determines the degree of 
resistance that the sound waves encounter. The amount of cohesion between molecules that constitute the 
different tissues results in differing degrees of resistance to the passage of the ultrasound. For this reason, 
most of the ultrasound waves that encounter fibrous tissues are reflected and not transmitted to deeper 
regions, while those that propagate through liquid structures, like a cyst or blood vessel, are barely reflected. 
It should be emphasized, however, that substances with too low a density, such as air, also fail to propagate 
ultrasound waves because there are too few molecules to propagate the sound wave. The difference in the 
acoustic impedance of various objects forms the essence of clinical ultrasound, as it allows the definition of 
acoustic interfaces that variably reflect ultrasound waves. These interfaces represent the various tissues 
through which the sound wave passes. 
While passing from tissues of low to high acoustic impedance, the ultrasound waves are modified by the 
angle of contact with the tissue interface and the incident surface. As long as the angle of incidence is equal to 
90°, acoustic interfaces having a smooth surface allow almost complete (specular) reflection of the ultrasound 
waves. Ultrasound waves that hit objects having a high acoustic impedance at an angle of 90° result in echoes 
that return to the transmission source, allowing the operator to establish the exact depth of the reflecting 
structure. That depth is equal to the product of the speed of sound waves through tissue (1542m/sec) and 
half of the time that elapses between their transmission and their reception (d=v*t/2). In situations where the 
angle of incidence is not exactly 90°, the ultrasound beam is partially reflected at an angle that equals the 
angle of incidence (i.e., a non-specular reflection) and partially refracted through the tissue. In these 
circumstances, the magnitude of the refractive deviation, caused by the marginally varied speed of propagation 
of the ultrasound waves through various tissues, is proportional to the differences of the acoustic impedance of 
the two tissues or objects (Figure 2.1). Incidentally, this forms the basis of many common acoustic artifacts 
discussed below. For a structure to either reflect or refract a sound wave, however, it must be at least as thick 
as a quarter of the length of the incident sound wave. Therefore a wave of 7.5 MHz can be reflected from 
structures > 0.038 mm thick, while a wave with a frequency of 2.5 MHz requires structures to be at least 
0.15mm thick to reflect or refract the wave. This explains why the spatial resolution of an ultrasound probe is 
directly proportional to its frequency, as will be detailed later. 
Click on the image to see a larger view 
Figure 2.1. Schematic representation of reflection and refraction of the ultrasound beam. In A the angle of incidence 
of the wave A with the surface of the tissue is less than 90°, resulting in a partial non-specular reflection (Al) and a 
partial transmission through the tissues with a specific angle of refraction (All). In B, the angle of incidence between 
the sound wave B and the tissue surface is equal to 90°. Therefore, the sound is partially reflected in a specular 
fashion (Bl) and partially transmitted through the tissue without refraction (Bll). 
Structures that are smaller than the ultrasound wavelength and have rough surfaces produce non-specular 
reflections or acoustic dispersion. An ultrasound wave crossing cellular or connective-tissue interfaces results in 
the formation of infinite echoes that are reflected in various directions which, together, determine the 
characteristic echogenic properties of parenchyma and viscera. 
Reflection, refraction, dispersion and thermal absorption determine the fate of sound waves through tissues, 
and together, they define acoustic attenuation (i.e., the loss of energy from the sound wave). This attenuation 
is directly proportional to the frequency of the ultrasound waves and is largely a function of the interaction 
between sound waves and the ultrastructural components of the tissues (Figure 2.2). For this reason, high 
frequency ultrasound probes have smaller penetration depth compared to low frequency probes. This 
relationship between the frequency of ultrasound waves and their attenuation in various tissues is illustrated in 
Table 2.1. Probes are described by a unit of measurement called "half-power distance", which is the distance at 
which the power of a sound wave is 50% of the emitted power. As can be seen in Table 2.1, the half-power 
distance for sound waves propagating through air is very short. This is why it is often necessary shave hair and 
to apply acoustic gel close to the tissue of interest, both of which reduce the interposition of air between probe 
and patient which would otherwise inhibit the examination. 
Click on the image to see a larger view 
Figure 2.2. Schematic representation of the phenomenon 
of attenuation. The ultrasound beam A, passing through 
the object, collides with (and is reflected by) fewer 
particles than ultrasound beam B. This difference is 
determined by the wavelength. Over a similar distance, 
ultrasound beam B will be attenuated more than 
ultrasound beam A since the energy of the beam is 
reduced by every contact with the particles in the tissue - 
the more contacts with particles, the lower the depth of 
penetration. 
The ultrasound probes, or transducers, act as both transmitters and receivers of the ultrasound waves. Their 
"hearts" are made of piezoelectric crystals that, when subjected to an electrical current, become deformed, 
producing sound waves of a specific frequency. Upon reflection from tissue interfaces, the sound waves impact 
these same crystals, causing them to vibrate at a specific frequency, which in turn is converted into an 
electrical current with a potential difference (voltage) that is correlated with the number of the returning 
echoes. 
The transmission and reception of sound waves by the transducers do not occur simultaneously. Instead, 
after a brief transmission burst or "pulse" of sound waves (usually 2-3 cycle lengths), the transducers activate 
their reception mode and listen for reflected echoes. Thus, the transducer spends well over 99% of the time 
receiving and <1% of the time transmitting. The frequency of repetition of the pulses (PRF - Pulse Repetition 
Frequency) expresses the number of ultrasound pulses per second (and is therefore expressed in Hertz). The 
duration of each ultrasound pulse is inversely proportional to the operating frequency of the probe - the lower 
the probe frequency, the longer the pulse duration. The PRF is directly proportional to the probe frequency. 
This is important, since in order to obtain an accurate image, it is necessary that the transducer receives all the 
echoes generated by the first pulse prior to transmitting a second pulse. Otherwise, echoes from deep 
structures with a longer travel time, if received after transmitting the second pulse, would be misinterpreted as 
having been reflected faster than they really were and would be assigned a more superficial reflecting 
interface. This phenomenon, associated with some of the more common acoustic artifacts, can be reduced by 
decreasing the PRF when high-frequency probes are used to interrogate deep tissues. Reduction of PRF, 
however, necessarily reduces the temporal resolution of the image, since fewer pulses are transmitted. 
Table 2.1. Half-power distance of the ultrasound waves in objects of ultrasonographic 
interest. 
Tissue 
Half power distance (cm) 
2 MHz probe 5 MHz probe 
Water 380 54 
Blood 15 3 
Soft Tissues 1.5 0.5 
Muscle 0.75 0.3 
Bone 0.1 0.04 
Air 0.05 0.01 
Resolution is the ability to identify two points lying above and below each other or beside each other. It is 
therefore obvious that the higher the resolution of an image, the greater the detail of the structures being 
imaged. The important types of resolution in ultrasonography are axial (up-down), lateral (side-to-side) and 
temporal (sweep speed) resolution. 
Axial, or longitudinal resolution is the ability to distinguish two points positioned on the same line as the 
propagated incident ultrasound wave (i.e. points that lie above and below each other, relative to the 
transducer). Axial resolution is greatly influenced by the transducer frequency, such that, for two points to be 
distinguished from each other, they need to be separated by a distance that is at least half the wave-length of 
the transmitted pulse. Therefore, the axial resolution is directly proportional to the operating frequency of the 
probe (Figure 2.3). 
Click on the image to see a larger view 
Figure 2.3. Schematic representation of axial resolution. 
The ultrasound beam emitted from the 3MHz probe A has 
a wavelength that does not allow visualization of particle 
2. The axial resolution for probe B probe, with a 
frequency of 6 MHz allows visualization of particle 2, 
because the wavelength is halved, and therefore collides 
with and is reflected from particle 2. 
Lateral resolution is the ability to distinguish two points at the same distance from the ultrasound source, but 
on different longitudinal or radial axes. In practical terms, it is the relative resolution of points along the same 
circumference, perpendicular to the longitudinal axis (Figure2.4). In order to best understand this type of 
resolution in the context of an ultrasound image, which is the product of multiple adjacent ultrasound beams, it 
is important to realize that these beams, while passing through tissues, diverge radially as their distance from 
the transducer increases. This results in the formation of the so-called "near" and "far" fields within the 
ultrasound image. The near field is the area interposed between the transducer and the point at which the 
ultrasound beams begin to diverge; beyond this point is the far field. Obviously, the lateral resolution within 
the near field is better than that within the far field (Figure 2.5). The magnitude of the divergence of the 
ultrasound beams in the far field is inversely proportional to both the transmission frequency and the width of 
the transmitting surface of the probe. Thus, a high-frequency probe has a better lateral resolution than a low-
frequency probe, and a linear probe has a much better lateral resolution and a bigger near field than a sector 
probe or a curvilinear probe. The point at which the ultrasound beams begin to diverge is known as the focal 
point and is considered the point at which the optimal lateral resolution is obtained. On many machines, this is 
adjustable or multiple focal points can exist, allowing maximal lateral resolution at several near field locations 
or depths (Figure 2.6). 
Click on the image to see a larger view 
 
 
Figure 2.4. Schematic representation of lateral resolution. (A) Particles 1, 3, 5 and 7, which lie in the image field are 
not visualized, and therefore do not contribute to the formation of the echocardiographic image. (B) Increasing the 
number of ultrasound beams of appropriate wavelength results in all 7 particles in the image field being visualized, 
and therefore contributing to the image. 
 
 
Figure 2.5. Schematic representation of the decrease in 
lateral resolution with increasing depth of imaging. In the 
near-field sectors, the ultrasound waves are effectively 
parallel and in close proximity to each other, allowing 
them to intercept all the particles in the field (A); 
however, in the far-field sectors, the waves diverge, so 
that some of the particles that are similarly spaced as 
those in the near-field sector cannot be visualized (B). 
Figure 2.6. Schematic representation of the variation of 
lateral resolution by means of focusing the ultrasound 
beam. Ultrasound waves above and at the focal point 
(indicated by the arrow) are nearly parallel, thereby 
improving the lateral resolution in these sectors. Distant 
to the focal point the waves diverge, reducing lateral 
resolution. 
Temporal resolution is the capacity to refresh the images visualized on the screen such that the structures on 
the screen change with changes in shape and position of the structures being imaged. Temporal resolution is 
generally defined by the frame rate. A way of understanding the concept of temporal resolution is to imagine 
early cinematography, where images were recorded with mechanical cameras that acquired only a few 
photographs per second. When these films were subsequently played, the movements of objects and people 
appeared jerky and discontinuous. The slow sampling rate of the early cameras resulted in phases of 
movement that were either captured as photos or were lost, penalizing the fluidity of movement in playback. 
In the sub-section about technical characteristics of transducers, the importance of temporal resolution in 
echocardiography will be detailed with methods of changing the frame rate to optimize the image. 
Box 2.1 summarizes the main concepts of the physics of ultrasound, as utilized in clinical echocardiography. 
Box 2.1 
Ultrasound. Sound waves with frequencies greater than 20 kHz. For diagnostic purposes, 
ultrasound frequencies range from 2 to >10 MHz. The speed of ultrasound waves in tissues is 
assumed to be constant at 1542 m/sec. The relationship between speed, frequency and wavelength 
is represented by the following equation: 
speed (m/sec) = frequency (Hz) x wavelength (m) 
If speed is constant, the frequency is necessarily inversely proportional to the wavelength. The 
term ultrasound beam identifies a series of ultrasound waves that propagate through an object in a 
single direction and are transmitted at the same time from the same source. The image field (image 
sector) is defined by a series of ultrasound beams of negligible thickness arranged side-by-side, and 
is typified by two-dimensional echocardiography. 
Acoustic impedance is the product of the density of tissue and the speed of sound in the tissue: 
acoustic impedance (z) = speed (v) x tissue density (ρ) 
However, since the speed of sound in tissue is constant, the acoustic impedance depends 
exclusively on the density of the tissue. From the difference between acoustic impedance of two 
adjacent tissues, the percentage of reflected and transmitted sound at the tissue interface can be 
determined as the sound wave passes from one tissue to the other. The amount of the reflected 
echo is directly proportional to the difference in acoustic impedance between the two tissues. 
Generally, the differences between the acoustic impedance of animal tissues are minimal allowing 
most sound waves to penetrate through the tissue interface. This characteristic is beneficial in 
clinical ultrasonography, because the beam is not entirely reflected from an interface, but is largely 
transmitted beyond the interface, and therefore able to reflect off deeper interfaces, allowing the 
operator to visualize structures at multiple depths. Bone and gas have very high and very low 
acoustic impedance, respectively. Consequently, the ultrasound beam, when it meets a tissue-bone 
or tissue-gas interface, is virtually completely reflected and is therefore unable to reflect off deeper 
structures. Therefore, when imaging, it is necessary to choose an acoustic window that avoids 
placing bone or gas between probe and organ being visualized. The same reasoning lies behind the 
use of an acoustic coupling gel applied to the skin of the patient so that the sound waves avoid the 
interposition of air between probe and skin. 
Reflection. When an ultrasound beam meets an interface of a tissue that is smooth and 
perpendicular to the direction of propagation of the ultrasound, with dimensions comparable to the 
wavelength of the ultrasound wave, the phenomenon of the specular reflection occurs. 
Dispersion. The dispersion of the ultrasound beam occurs when the sound wave meets a series of 
small and irregular interfaces in the parenchyma of an organ. The term of non-specular or diffuse 
reflection is also used. Dispersion is independent of the angle of incidence of the beam. Many small 
echoes are formed that, cumulatively, become visible. These echoes are responsible for the 
"characteristic architecture" of the parenchyma of many organs. The dispersion increases with 
transducer frequency. 
Absorption. Refers to the conversion of the mechanical energy of the sound wave into thermal 
energy. Heat is produced by friction between tissue molecules which vibrate longitudinally at the 
same frequency as the ultrasound wave. 
Attenuation.Is the loss of power of an ultrasound beam either before reaching a tissue interface, 
where it would be reflected, as well as after reflection during the return of the reflected wave to the 
transducer, such that some of the sound waves constituting the beam fail to reach the transducer. 
The attenuation is directly proportional to the frequency of the ultrasound wave and is affected by 
absorption, reflection and dispersion of the ultrasound beam. Distal to structures that cause a high 
degree of attenuation, areas lacking echoes are produced (shadows), while distal to structures that 
cause a low degree of attenuation, stronglyechogenic areas are produced (enhancement). 
PRF or pulse repetition frequency. The ultrasound waves used in clinical ultrasonography occur 
in pulses, emitted as salvoes or pulses of 2 or 3 wavelengths. After transmitting the pulse, the 
transducer then listens for returning echoes. In order to accurately interpret the distance between 
the probe and imaged structures, all of the echoes generated by a pulse must be received before 
transmitting the next pulse. The time between successive pulses is known as the PRF. Lower 
frequency probes with greater penetrating depth of the ultrasound beam have a lower PRF to allow 
sufficient time between pulses for all reflected sound waves to be detected. 
Resolution of the image. Refers to the ability to distinguish two points, or the position of a single 
point over time. The greater the resolution the smaller the distance between the two points that 
can be distinguished. 
 Axial resolution refers to the ability to distinguish two points along the longitudinal axis of 
the ultrasound beam (one above the other). This is determined by the operating frequency 
of the transducer - high frequencies, which have short wavelengths, can distinguish 
reflections of structures that are extremely close together. 
 Lateral resolution refers to the ability to distinguish two adjacent points perpendicular to 
the longitudinal axis of the ultrasound beam. This depends on the degree of divergence of 
the ultrasound waves that constitute the beam which, in turn, is influenced by the 
dimensions and shape of the transducer surface, and the distance from the transducer. 
The shape and dimensions of the transducer surface determine the number of parallel 
beams that can be emitted, which in turn affects the lateral resolution. The final factor that 
affects lateral resolution is the focusing of the ultrasound waves to a focal point where the 
resolution is maximal. 
Since axial resolution is generally higher than lateral resolution, due to greater flexibility of 
frequency than of probe design, it is advisable that all the measurements possible should be 
obtained along the longitudinal axis of the ultrasound beam. In an echocardiographic image, it is 
possible to recognize areas with different resolutions. 
 Fresnel Zone: Very close to the probe, this zone has complex interference phenomena that 
makes it difficult, if not impossible, to distinguish structures. 
 Fraunhofer Zone: This zone lies beyond the focal zone, where ultrasound waves diverge 
quickly at the expense of resolution. This is effectively the "far-field". 
 Focal Zone: This zone is the part of the image field where the resolution of the images is 
optimal. It surrounds the focal point by several centimeters and represents the area where 
the ultrasound beam is narrowest. 
 
SEARCH RESULT #: 2 TITLE: Formation of the Image; Artifacts AUTHOR(S): 
ADDRESS (URL): http://www.vin.com/Members/Proceedings/Proceedings.plx?CID=ECHO2007&PID=20490&O=VIN 
 
FORMATION OF THE IMAGE 
Two-dimensional and M-mode echocardiography use identical fundamental principles of ultrasound. The 
electrical circuits of the ultrasound system supply pulses to one or more piezoelectric crystals within the 
transducer head which function both as transmitters and receivers of ultrasound waves. During the 
transmission phase, the electronic circuit generates a short discharge (from 500 to 1500 pulses/sec) of 
alternating current, which vibrates the piezoelectric crystals at a specific frequency, creating the ultrasound 
pulse. This pulse, penetrating through tissues, meets acoustic interfaces which create ultrasonic reflections 
along small planes with the number of reflected waves being proportional to the difference of acoustic 
impedance between the objects. The returning echoes hit the piezoelectric crystals within the transducer, 
causing them to vibrate at a specific frequency. The strength of the vibration and the number of vibrating 
crystals create an electrical current with a potential difference proportional to the intensity of the echoes. A 
gray scale is then assigned for every potential difference (with white being most intense, and black 
representing a complete absence of echoes) with an image displayed on the screen similar to that displayed in 
Figure 2.7. In order to place a particular point at a specific location on the screen, the ultrasound system first 
calculates the distance of the source of the echoes from the transducer by means of the formula D = V x T/2, 
where T= the time from transmission to reception of the sound wave at the transducer and V = the velocity of 
sound through tissues (1542m/s). 
Click on the image to see a larger view 
Figure 2.7. Schematic representation of the formation of 
the image by the echocardiographic machine. Structures 
with different acoustic impedance are represented on the 
monitor by various shades of gray. In this simplistic 
cartoon, the outline of the bony tissues is represented on 
the screen as a white image (indicating highly echogenic 
tissues), even though in reality all the ultrasound waves 
would be reflected from the proximal periosteal surface. 
ARTIFACTS 
The ultrasound machine can also produce images that are not accurate, or do not truly represent anatomical 
structures, but rather are the consequences of exaggerations or limitations of physical characteristics of 
ultrasound waves that result from interactions with specific normal or pathological structures. 
The echocardiographic image, therefore, must be examined and interpreted in its entirety, both in terms of 
morphology and function as well as "echo-structurally" (i.e., with consideration to the acoustic patterns 
characteristic of the tissues being imaged), always considering whether reflections are real or artifactual. 
Reverberation 
Reverberation artifact results in multiple perfect (specular) reflections, caused by two highly reflecting 
interfaces, an example of which is the skin-probe interface. In echocardiographic images, reverberation 
appears as numerous equidistant, parallel hyperechogenic curved lines, beginning at the probe-skin interface 
(top of the image) and extending some distance into the image. Reverberation artifact occurs when the 
reflected echo is intense and, upon reaching the transducer, it interacts with the crystal. Part of the energy of 
the echo is transformed into an electrical current and is recorded as a luminous point on the screen, while the 
other part is reflected back into the tissue and travels to the same interface that had created the original 
reflection, is reflected from this interface a second time and returns to the transducer. Since the time to the 
second detection is double that of the first, the second reflection is denoted on the screen as a second point at 
twice the depth of the first. This repeated reflection can occur several times, resulting in a series of ever-
deepening points on the screen (Figure 2.8). 
Click on the image to see a larger view 
Figure 2.8. Reverberation artifact: This appears as 
multiple highly echogenic evenly spaced arcs of increasing 
circumference. 
Mirror Effect 
Mirror effect is a particular type of reverberation artifact produced by an interface of moderately high 
reflectivity (e.g. the pericardium) that is less intense than those producing typical reverberation artifacts. In 
practical terms, the artifact usually appears deep to the white line that is causing the relatively intense 
reflection (e.g. a pericardial reflection) as a repetition of some particular structure or organ below the actual 
location of the true structure (Figure 2.9). 
Click on the image to see a larger view 
Figure 2.9. Mirror Artifact. The left ventricle (A) and left 
atrium (B) are "mirrored" in the lower part of the image 
(AI and BI), giving rise to a false image separated from 
the real oneby a white hyperechoic line (the 
pericardium). 
Spontaneous Echocardiographic Contrast 
Spontaneous echocardiographic contrast (smoke) is visualized as the appearance of small areas of medium 
echogenicity that move within the cardiac chambers or vessels, consistent with the flow of blood. The origin of 
this artifact is generally attributed to microaggregates of red blood cells that tend to occur with 
hypercoagulable states or low flow conditions (stasis). Spontaneous echo contrast is also a normal 
phenomenon in the horse, and less frequently in the dog. It is considered abnormal in cats (Figure 2.10). The 
ability to see spontaneous echo contrast is somewhat dependent on the transducer frequency, and has been 
reported in mice using ultra-high-frequency transducers (50MHz). 
Click on the image to see a larger view 
Figure 2.10. Spontaneous echocardiographic contrast: 
appears as hyperechoic points within the cardiac 
chambers that, in real time, move and swirl with blood 
flow. 
Side Lobe Artifacts 
These are diffuse reflections that originate lateral to the structures that actually reflect the beam. The physics 
behind side-lobe artifacts are relatively complex. Echocardiographically, they appear as weak gray curvilinear 
objects just lateral to the central echocardiographic field (Figure 2.11). Because these artifacts are extremely 
weak, they are generally only visible when the central beam penetrates liquid (hypoechoic) structures such as 
cardiac chambers. This type of artifact is often visible within the left atrium, where reflections of the ventricular 
endocardium beside the left atrium are visualized within the left atrium (Figure 2.11). The artifact can be 
eliminated by optimizing the image or centering the hypoechoic (fluid-filled) structure on the image, which 
reduces the number of ultrasound rays that hit the lateral parenchyma. 
Click on the image to see a larger view 
Figure 2.11. Side lobe artifact. This is caused by the 
erroneous positioning of echoes originating lateral to the 
site of the artifact. These artifacts usually become visible 
within cardiac chambers (which are hypoechoic and are 
positioned beside hyperechoic structures, such as the 
pericardium). 
Acoustic Enhancement 
This phenomenon occurs distal to poorly-reflective tissues, mostly collections of fluid (e.g., cysts), where areas 
directly distal to the fluid-filled cavity are hit by ultrasound waves of greater intensity than adjacent areas that 
do not have fluid cavities proximal to them. The waves that have passed through the proximal fluid region are 
attenuated minimally by the fluid. Therefore, the areas distal to the fluid cavities appear brighter than the 
adjacent areas (Figure 2.12). 
Click on the image to see a larger view 
Figure 2.12. Distal enhancement. The structures 
indicated by the arrows appear highly echogenic, when 
the ultrasound waves meet an acoustic interface between 
tissues of very different acoustic impedance. Generally, 
distal enhancement occurs deep to hypoechoic structures, 
which attenuate few sound waves - this allows the tissue 
interface immediately distal to the hypoechoic structure to 
reflect more sound waves than neighboring tissue. 
Acoustic Shadows 
These are zones on the echocardiographic image where the reflections are very weak or totally absent. The 
shadow zone appears, therefore, as an "asonic" or dark area of variable shape which can be distinct or can 
fade. The artifact occurs as a result of failure of reflection of any sound waves because of attenuation of the 
waves in tissues proximal to the area. As such, these artifacts occur immediately distal to tissues with very 
high acoustic impedance, which inhibit the passage of any ultrasound waves. Typically, acoustic shadows are 
caused by bones, calculi or gas (Figure 2.13). Gaseous shadows differ from bone/calcium shadows. 
Click on the image to see a larger view 
Figure 2.13. Acoustic shadow artifact: in this example, 
the shadow artifacts appear as two vague and indistinct 
triangular areas devoid of echoes, because the ultrasound 
waves have been totally reflected by two highly reflecting 
structures (ribs). 
 
 
ECHOCARDIOGRAPHIC EQUIPMENT 
Echocardiographic equipment essentially comprises the monitor, the body of the machine (including the 
keyboard, loudspeakers, computer, and a frame with wheels) and probes or transducers (Figure 2.14). (Please 
refer to the section on echocardiographic probes for more detail about these.) 
Click on the image to see a larger view 
Figure 2.14. Ultrasound machine of the type used in 
veterinary teaching hospitals. The unit has been adapted 
to allow portability in a hospital environment where the 
machine is used for imaging both small and large animals. 
The echocardiographic unit is usually connected to a printer, video recorder or digital image storage (most 
common on new machines) for recording of images and movies, allowing permanent documentation of the 
examination and subsequent post-hoc analysis. Many machines also allow the operator to perform 
measurements and calculations on pre-recorded images at some time after the actual examination (off-line 
analysis). Newer systems also provide stand-alone imaging software for image analysis at a computer separate 
from the ultrasound machine. Some portable echocardiographic units, like the one in Figure 2.15, are equipped 
with LCD screens, substantially reducing the dimensions of the system and increasing portability. Occasionally, 
these units can be connected to a traditional screen for better visualization. All machines come with the ability 
to attach an electrocardiograph (Figure 2.16) which permits recording of a synchronized ECG tracing with the 
echocardiographic image. Synchronous ECG recording allows the user to accurately identify systolic or diastolic 
frames. This is important for evaluating timing of certain events and specific indices such as the pre-ejection 
period. It is considered good echocardiographic practice to connect the ECG whenever performing an 
examination. In some situations, experienced echographers may omit the ECG. However, this can limit the 
value of the diagnostic information obtained during the examination. 
Click on the image to see a larger view 
 
 
Figure 2.15. Portable ultrasound machine, equipped 
with an LCD display in order to reduce the size and 
weight. 
Figure 2.16. Portable ultrasound machine mounted on a 
cart and connected to cables for the simultaneous recording 
of the electrocardiogram. 
Echocardiographic Controls 
The controls present on most machines include: 
 Those that regulate the dimensions and the position of the image; 
 Those that influence the quality and the intensity of the image. 
The depth control regulates the distal limit of the ultrasonic field and therefore the relative sizes of the 
imaged organs. It does not regulate the depth to which the ultrasound beam travels, but only the depth to 
which reflected waves are processed (recall the relationship between velocity, time and depth). The lesser the 
depth, the larger the structures in the near-field will appear. It is important to understand that reducing the 
depth of field, even if it offers better recognition of fine details (e.g., fine irregularities of a surface that would 
have otherwise appeared smooth, or a high-frequency small amplitude vibration of a valve), will reduce the 
overall image quality by increasing graininess and pixilation. The reference depth scale always appears to the 
left or right of the image and is automatically adjusted with changes in depth and allows experienced operators 
to estimate the size of structures or organs without specific measurement (Figure 2.17). Usually the scale is in 
0.5cm or 1cm increments, often with highlighted 5cm increments. 
Click on the image to see alarger view 
Figure 2.17. Portions of the ultrasound screen. On the 
left of the ultrasound image is the grey scale, and the 
depth scale, in centimeters. The small triangle next to the 
scale is the focal point, which can be positioned by the 
operator. Next to the apex of the image field, (in this case 
to the right of the apex) is another triangle indicating the 
position of the probe in relation to the image. 
The other scale of reference on an ultrasound image appears at the bottom of the screen when imaging in 
either Doppler mode or M-mode, and identifies the timing or speed of the image scrolling (Figure 2.18). It 
generally denotes 1 second intervals (or fractions of a second). Speed controls allow the operator to alter the 
"sweep speed", so as to provide greater or lesser detail to temporal events. With standard imaging, a 
moderate sweep speed is used and is increased when examining fine details of the oscillations of various 
structures, when imaging subjects with high heart rates (Figure 2.19), or when obtaining extremely precise 
temporal measurements. 
Click on the image to see a larger view 
 
Figure 2.18. Schematic representation of the monodimensional (M-mode) diagram of an oscillating point (producing a 
sine wave). In A, the sweep speed of the image is fast so that approximately 3.5 sec are displayed on the screen. In 
B, the sweep speed is much slower resulting in 7 sec being displayed on the screen. On the left of the diagrams is the 
depth scale. 
Figure 2.19. M-mode image of the relatively fast cardiac 
activity recorded in a cat. The sweep speed of the image 
(3 sec per screen) is set up to obtain a suitable temporal 
resolution for the differentiation of several cardiac cycles 
(arrows). 
Many systems allow regional amplification of the image, via the so-called "zoom" function, which can also be 
done off-line in the more sophisticated machines. Expert operators often zoom in on specific regions or 
structures when performing 2-dimensional imaging, preferring smaller but more detailed views of specific 
structures and their movements. Using the zoom function takes practice, because the general references that 
the operator uses to optimize the image or to identify what is being examined, are often lacking. 
All systems have a sector size control which allows the operator to control the angular dimension of the 
image field (effectively, the size of the image slice), allowing the operator to reduce or increase the field 
visualized on the screen (Figure 2.20). Generally, the more the visualized area is reduced (the smaller the 
slice), the better the performance of the ultrasound machine because of reduced image processing. 
Specifically, the frame rate improves when the visualized area is reduced, which in turn allows improved 
visualization of rapidly moving structures. Therefore, in echocardiography, it is advised to use the smallest 
effective sector. This is especially important for color Doppler imaging. 
Click on the image to see a larger view 
 
 
Figure 2.20. Image fields of different width. In echocardiography, as narrow a field as possible is preferred (B or C), 
in order to optimize the frame rate and the reproduction of the movements of cardiac structures. In fact, the larger 
the field of view, the lower the number of images that can be acquired every second (in this panel, the frame rates 
vary from 62 frames/sec in A to 83 frames/sec in B and 127 frames/sec in C). 
Note: Imm/S = frame rate (frames/sec) 
Additional controls allow flipping of an image in either a horizontal or vertical plane to allow standard 
positioning of the image. 
Most of the controls found on an echocardiographic unit regulate the quality and the intensity of the image. 
Total input power (or gain) can be controlled on all units to increase or reduce the intensity of the reflected 
sound waves. However, because the intensity of the ultrasound waves decreases as they pass through the 
body, the echoes returning from the far field are weaker than those returning from the near field. Without 
correcting for this discrepancy, deeper structures of similar physical characteristics, or an organ that occupies 
both near and far fields, would appear to be heterogeneous in echogenicity with the more distant regions 
appearing less echoic than proximal regions. In order to overcome this problem, all echocardiographic units 
have a series of controls that allow reduction of the gain of near-field echoes and intensification of far field 
echoes called Time Gain Compensation (TGC) controls, that affect gain within very specific depths. Using TGCs 
(Figure 2.21), it is possible to adjust the image observed on the screen, increasing or reducing the 
amplification (gain) of selected regions. Expert ultrasonographers adjust the TGCs for every patient and 
frequently readjust throughout the examination to provide an optimal image. 
Click on the image to see a larger view 
Figure 2.21. Ultrasound machine console: 1) Time-Gain 
Control potentiometers for regulating gain at different 
depths (2) Track ball, analogous to a computer mouse, 
for moving cursors on the screen (3) Controls for 
selecting different ultrasound applications (2D, M-mode, 
Doppler etc); (4) Image freeze button (5) Control for 
adjusting overall image gain; (6) Controls for adjusting 
Doppler and Color Doppler gain; (7) Control for angle 
correction of pulsed-wave Doppler beams; (8) 
Alphanumeric keyboard; (9) Control panel for adjusting 
depth, focus, frequency, and image size (zoom); (10) 
Control panel for adjusting Doppler settings; (11) 2D-hold 
button for providing reference images during spectral 
Doppler interrogation; (12) Control for setting the 2D 
update rate during spectral Doppler imaging; (13) 
Spectral Doppler baseline adjustment; (14) Image 
capture/store/record button; (15) Control panel for 
selecting imaging pre-sets, probes, and patient 
information, retrieving archived images, and setting up 
default pre- and post-processing settings; (16) Control 
panel for performing measurements and calculations of 
ultrasound images. 
Some low-end units only have controls for adjusting the total power (or total gain), the mid-field gain and far-
field gain. A good rule-of-thumb for setting the gain is to begin every examination with the total gain set at 
approximately 50%, the near-field gain around 25% and the far-field gain at 75%. This usually results in 
reasonable image and allows the operator flexibility in additional adjustments. 
Echocardiographic units have additional power controls, which alter the output power of the ultrasound 
waves, as compared to gain controls which alter the amplification of the echoes received by the transducer 
(input power). Some machines allow the amplification of weak signals and the attenuation of more intense 
signals. Output power is usually set at 100% unless performing fetal examinations. 
Additional controls are available for sensitivity, attenuation, edge enhancement and rejection. Sensitivity 
controls are filters that suppress low intensity reflected signals. These low-intensity reflections are usually 
responsible for creating slight graying of areas that should appear black. Therefore, eliminating these signals 
improves the "blackness" of anechoic structures. Conversely, attenuation controls filter high-intensity signals 
thereby reducing the impact of hyperechoic structures on image quality. Edge enhancement filters allow 
greater contrast between interfaces. 
Increasing the output power increases the density of the transmitted ultrasound waves, increasing the image 
brightness, and allowing visualization of structures with weaker reflections. However, if the gain is too high, the 
brightness of the more echoic structures may drown out echoes of the less echoic structures. A fundamental 
concept for the correctinterpretation of images is that every image is the result of the application of a specific 
"code" which assigns a specific level of gray to each specific intensity of the reflected ultrasound signal. 
Intensity of reflections is displayed along a gray scale (from white to black). Additionally, it is possible to use 
various scaling to alter the image, providing high-contrast or low-contrast images, by adjusting the "gray scale 
gamma curve" (Figure 2.22). Most of these are pre-programmed, and allow the operator to select an overall 
gray scale range, rather than doing manual adjustments. For optimal echocardiographic visualization, it is 
always preferable to select range curves that provide strong differences between tissues, such as valves, 
chordae tendineae, serosa, endocardium, and anechoic regions, such as cardiac chambers or vessels (Figure 
2.23). In other words, use range curves providing a high-contrast image with a broad but limited gray scale 
range. However, selecting gray scales that provide excessive contrast diminish the more subtle differences 
within tissues such as myocardium and may reduce the diagnostic quality of the image. The pre-programmed 
gray scale options vary in number between units. Additionally, colorized scales are also available, and these 
can offer advantages in certain situations, especially when imaging pericardial effusions, because the human 
eye can discriminate between colors better than grays. 
Click on the image to see a larger view 
 
Figure 2.22. Two dimensional echocardiographic image with different range curves: In panel A, a broad linear grey-
scale range has been utilized; in panel B the gray-scale curve accentuates the grays. 
 
 
Figure 2.23. Two dimensional echocardiographic image (A) with a high-contrast sigmoidal grey-scale curve (B) that 
accentuates the differences between the hypoechoic and hyperechoic structures. 
Other controls alter image quality by including or excluding reflected ultrasound pulses of specific intensities. 
One of these is the DNF (Dynamic Noise Filter), which eliminates some frequencies that cause image flutter on 
the monitor. This filter, installed on machines with mechanical sector probes, should be inactivated during an 
echocardiographic examination as it affects the imaging of valve motion. Similarly, persistence filters, which 
"smooth" the image by slightly overlapping successive images, should be turned off during echocardiographic 
examinations so that fine and rapid motion can be detected. 
The speed of acquisition of the images from the transducer (FR - Frame Rate) can be modified by the 
operator to some extent. In general with echocardiography, image quality is enhanced by high frame rates. 
The FR can usually be identified on the screen as a frequency (frames/sec) (Figure 2.20: Note: Imm/S = frame 
rate (frames/sec)). 
In abdominal ultrasonography, where relatively immobile organs are studied, a low FR does not create a 
problem; indeed, it can help to reduce the effect of small vibrations caused by oscillation of the transducer 
head. On the other hand, in echocardiography, a higher FR improves visualization of structures that move at a 
relatively high speed, such as valves. Therefore, with echocardiography the FR is most important factor 
affecting the generation of an accurate image and the operator must always attempt to use settings that favor 
a high FR. Another fundamental difference between systems designed for cardiology vs general imaging is the 
number of focal points. Since the FR is inversely proportional to the number of focal points, the FR in 
echocardiographic units utilizes a single focus, which is then manually positioned during the examination to 
optimize the region being interrogated. Generally, the focal point is positioned towards the distal end of the 
cardiac image, since image resolution proximal to the focal point does not suffer as much as that distal to the 
focal point. 
Box 2.2 summarizes some features of echocardiographic equipment that affect the rate of image acquisition. 
Box 2.3 summarizes the features of controls that influence the quality of images. 
Box 2.2 
Frame rate - this is the number of images formed per second. This is dependent on the maximum 
depth of the returning echoes being processed. The greater the depth, the longer the time 
necessary for the return of the echoes to the transducer and, consequently, the lower the FR. The 
wider the image sector, the lower the FR. The more focal points, the lower the FR. Thus, in order to 
maximize FR, the image field should be narrow, shallow and unifocal. 
Frame averaging - This averages the frames, resulting in a reduction of the differences between 
the various images, reducing the graininess of the image, but reducing edge resolution. The FR and 
frame averaging are inversely related. 
Box 2.3 
Rejection levels - These set the limits for eliminating weaker echoes (which do not contribute in a 
meaningful way to the formation of the image) from all the depths. 
Gain control - This amplifies all the returning echoes in uniform way, regardless of the depth at 
which they were reflected. 
Time gain compensation - These controls amplify specific regions of the ultrasound image (i.e., 
only reflected echoes with specific time signatures). These controls allow the operator to 
compensate for the natural attenuation of echoes from distal fields. 
Gray scale - Two-dimensional images are portrayed on the screen along a gray scale. In situations 
where it is important to distinguish differences between low-intensity echoes or between high-
intensity echoes, more levels of gray are assigned (low contrast). In situations where it is more 
important to distinguish differences between high-intensity and low-intensity echoes, the gray scale 
is reduced (high contrast). In echocardiography, where both of these situations are important, the 
gray scale has a sigmoidal shape (Figure 2.23), which provides a relatively well-contrasted image 
with a broad gray scale range. 
Power control - This modifies the voltage applied to the piezoelectric crystals to generate the 
pulse. Increasing the voltage increases of intensity of the ultrasound beam and, consequently, of 
the returning echoes, producing a brighter image. 
All ultrasound machines have dedicated computer hardware and software designed for image processing, 
display and storage. In addition to improving and displaying the images, the digital processor also allows the 
operator to perform various measurements and to store images in a buffer from which a specific image can be 
selected to be studied in detail. The stored and saved digital data (which today can occupy gigabytes of 
memory per study) can be re-examined remotely at a later time. Therefore, modern echocardiographic units 
are increasingly more and more powerful computers with immense real-time processing capabilities for 
generating high-quality images, measurement packages dedicated to the various imaging applications (2-
dimensional, M-mode and Doppler) and large drives for archiving of studies. 
 
SEARCH RESULT #: 2TITLE: Transducers AUTHOR(S): ADDRESS (URL): 
http://www.vin.com/Members/Proceedings/Proceedings.plx?CID=ECHO2007&PID=20492&O=VIN
 
Ultrasound waves are created by the piezoelectric effect, which is an example of the transformation of an 
electric field applied to a particular material (e.g. quartz or disc of ceramic material) into an elastic sound 
wave. This piezoelectric effect is reversible: the transducer not only produces ultrasound waves, but also 
detects and decodes the reflections (echoes) of these waves and converts them into an electrical current. With 
ultrasound transducers, many crystals are arranged along a surface, effectively constituting many point 
sources of ultrasound waves, which, when summed, produce a wave front or beam.Today, transducers are mostly electronic rather than mechanical. Electronic (phased-array) transducers are 
composed from tightly packed synthetic crystals arranged on a flat or curved surface such that the ultrasound 
waves emitted by each crystal remain near-parallel as they progress from the probe surface to the tissues 
being imaged. If the piezoelectric crystals are arranged on a flat surface (as with linear probes, such as the one 
depicted in Figure 2.24), the shape of the ultrasound field is rectangular (Figure 2.25). This shape limits the 
possibility of interrogating organs "hidden" behind other structures of the body, as in thoracic imaging where 
the ribs and lungs are positioned between the probe surface and the heart. However, because of the parallel 
nature of the ultrasound waves with linear probes, lateral resolution is maximal. To overcome the issue of 
imaging "around" interfering structures, such as bones or air, some probes have a small surface with a 
diverging beam (sector scanner) that can "squeeze" between interfering structures, at a cost of resolution. A 
curvilinear probe, or microconvex probe is a compromise between linear and sector probes. 
Click on the image to see a larger view 
 
Figure 2.24. A linear ultrasound probe. Figure 2.25. The rectangular ultrasound field that is obtained with 
a probe whose crystals are arranged linearly (linear probe from 
Figure 2.24). 
The probe characteristic that determines the imaging depth is the transmission frequency: the lower the 
frequency, the greater the imaging depth. The higher the probe frequency, the lower the maximum power that 
can be used to produce ultrasound waves, which leads to increased attenuation of the beam. Therefore, for 
feline echocardiography where image depth is usually <6 cm, probes with frequencies >7.5 MHz are generally 
used, although some feline patients may require lower-frequency probes to obtain an image. Additionally, 
Doppler imaging is often improved with lower-frequency probes than would be used for the same animal with 
2D imaging. Optimal probe selection is determined case-by-case and species-by-species (5 vs 3.5 vs 2.25 
MHz). Interestingly, depth limitations for low-frequency probes (2-3 MHz) are largely assigned by 
manufacturers, usually to a maximal depth of 30 cm, since greater depths are not required for imaging human 
patients. This can affect equine examinations, where cardiac structures may lie more than 30 cm from the 
probe. 
These days, many probes are multifrequency probes, able to generate several different frequencies. This is 
controlled digitally and provides a benefit to veterinary ultrasonographers because it reduces the number of 
probes that need to be purchased in order to cover the range of animal sizes to be examined. However, with 
multifrequency probes, the frequency that offers the best performance is the central frequency so the other 
frequencies are often underutilized. 
The anatomical and topographical objects that impede the 2-dimensional ultrasound beam through narrow 
acoustic windows used in echocardiography obviously do not represent a problem for a single linear ultrasound 
beam (effectively, an ultrasound wave). This feature is used in M-mode or Doppler imaging. Of course, a linear 
ultrasound beam cannot penetrate through bones or lungs; however, given that the thickness of the beam is 
negligible, these beams can find room through the narrowest of acoustic windows. 
Transducers used in echocardiography can be classified into 4 types based on the type of ultrasound beam 
transmission. 
Convex Transducers 
These represent a variation of linear transducers, and have crystals mounted in an arc to reduce the effective 
contact surface with the subject's body. The image field is consequently created by divergent ultrasound 
waves, slightly distorting the image. Convex transducers can be further classified as: 
 Convex, with a curvature surface > 20 millimeters; 
 Microconvex, with a curvature surface < 20 millimeters (Figure 2.26). 
These probes can be used for B-mode, M-mode and Doppler imaging (spectral and color). However, they are 
limited by relatively low frame rates. 
Click on the image to see a larger view 
Figure 2.26. An electronic microconvex probe with a 
frequency of 6.5 MHz. Note that the transmitted beam is 
relatively narrow because of the curving of the surface, 
and is comparable to the size of a wedding band. 
Mechanical Sector Transducers 
These transducers have crystals that generate ultrasound waves of a single frequency, which are rapidly 
oscillated or rotated by a mechanical motor. (In some probes, known as annular array probes, 2 arrays of 
crystals with different emission frequencies were used to double the frequency range.) Generally, in 
mechanical sector transducers, the movement is produced by a mechanical arm in the head of the probe 
connected to a motor. The angle of the image field is normally 90°, but it can also be wider or narrower. The 
reduced contact surface and the shape of the ultrasound beam make sector scanners the optimal type of probe 
for echocardiography (Figure 2.27). A disadvantage of mechanical transducers is their overall size (relatively 
large) and their delicate nature making them susceptible to damage if dropped or hit. These probes can be 
used for all echocardiographic applications. 
Click on the image to see a larger view 
Figure 2.27. Mechanical sector probes. The upper probe 
has a frequency of 3.5 MHz allowing imaging to 
approximately 30 cm. The lower probe, on the other hand 
has two piezoelectric crystals that produce ultrasound 
waves of different frequencies. This allows the operator to 
select multiple frequencies with the same probe - in this 
case 5 and 7.5 MHz. Note the markers on the probe that 
allow the sonographer to identify the orientation of the 
transducer by touch. 
Pencil Probes 
These are very specific transducers (Figure 2.28) and are the simplest form of transducer, emitting a single 
linear continuous ultrasound beam. They were the original probes used for M-mode imaging. Subsequently, M-
mode functionality was supplanted by continuous-wave Doppler. Thus, today, they are used solely for 
continuous-wave Doppler imaging as ancillary transducers on systems that do not have continuous-wave 
Doppler capability in the sector transducers (see chapter 3, Continuous-wave Doppler). Pencil probes provide 
very high-fidelity wave-wave Doppler images having only a single function and being low-frequency probes, 
and they are relatively cheap (~$2,000) but require substantial training to use. 
Click on the image to see a larger view 
Figure 2.28. A low-frequency "pencil" probe with a 
transmission frequency of 2 MHz. 
Phased Array Sector Transducers 
These transducers were developed with the evolution of digital electronics (Figure 2.29). The contact surface of 
the probe is relatively flat and small - generally smaller than mechanical sector transducers. The synthetic 
piezoelectric crystals are parallel to each other and activated in a very rapid regular sequence by software in 
the unit. The high cost of these transducers is sometimes an impediment to general veterinary use but they 
are becoming increasingly popular. Phased array probes can accommodate high frame rates and are used in 
all cardiac applications. The processing capability with many systems that offer phased-array probes allows for 
dual imaging (e.g. simultaneous real-time 2-dimensional and M-mode, or 2-dimensional and spectral Doppler) 
although at some expense to image quality. Very new probes can even provide two orthogonal 2-dimensional 
views simultaneously. In most systems currently used in veterinary medicine, pulsed-wave Doppler imaging is 
optimized by "inactivating" the simultaneous 2D imaging via "stand-by" mode to allow dedicated image 
processing tothe pulsed-wave signal. When using these probes for continuous wave Doppler, dual imaging is 
not possible. 
Click on the image to see a larger view 
Figure 2.29. A phased-array electronic sector probe, 
used in cardiology, with a frequency of 3.5 MHz. The low 
transmission frequency of this probe makes it most suited 
for use in larger animals. 
 
 
SEARCH RESULT #: 1 
TITLE: Techniques of Acquiring the Echocardiographic Image 
AUTHOR(S): 
ADDRESS (URL): http://www.vin.com/Members/Proceedings/Proceedings.plx?CID=ECHO2007&PID=20493&O=VIN 
 
ECHOCARDIOGRAPHIC IMAGING TECHNIQUES 
This section describes the techniques of performing B-mode and M-mode echocardiographic examinations. In 
the dog and cat, these techniques use the same acoustic windows and imaging planes as the Doppler 
techniques that are detailed later. In the horse, the imaging techniques are similar, but some points require 
additional explanation and are dealt with in the chapters on equine echocardiography. 
The graphical representations of M-mode and B-mode images are completely different from each other (Box 
2.4). Essentially, B-mode imaging displays cardiac anatomy and shape in realistic format. The M-mode image 
does not reproduce an anatomical reality, but produces a representation of the movements of a very specific 
cross-section of the heart over time. 
The M-mode image in Figure 2.30 represents, along the vertical axis, the depth of the region being imaged 
with the region closest to the transducer at the top and the region furthest away at the bottom of the image. 
Time is represented on the horizontal axis. Each tissue interface that generates an echo is represented by a 
single point. A serial acquisition of linear vertical images allows them to be stacked side-by-side, creating an 
image that displays the positions of these points relative to the transducer over time. Because these linear 
vertical images are acquired very rapidly, the change in position of these points is represented on the M-mode 
diagram as a line or small sine-wave. If the object approaches the transducer, the line moves up; if the object 
moves away from the transducer, the line moves down (Figure 2.30). If the movement is repeated in cyclical 
fashion, this will be obvious on the diagram. The degree of movement is represented by the amplitude of the 
deflection of the points. Movement perpendicular to the ultrasound beam cannot be detected by M-mode 
imaging. 
Click on the image to see a larger view 
Figure 2.30. Monodimensional diagram demonstrating 
movement of the heart, as it appears on the screen in 
real-time. The reference 2D scan is seen in the upper 
portion of the image, on which an arrow denotes the 
movement of the specific portions of the heart (two-
headed arrow, and single-headed arrow); the lower 
portion represents the M-mode image, with the arrows 
corresponding to the same movements of the same 
structures as seen in the 2D image. The sweep speed 
allows about 3 seconds to be displayed. T0 and T1 indicate 
two moments, in which the examined structures are in 
different positions at different points in the cardiac cycle 
(diastole and systole). 
Box 2.4 
A-mode (amplitude mode) ultrasonography is no longer used in diagnostic imaging (with the 
exception of ophthalmology) and is only of historical interest in echocardiography. An A-mode 
image somewhat resembles a spectrogram, with the echoes appearing on the monitor as spikes of 
different amplitude along a horizontal baseline. The height of the spike represents the intensity of 
the signal, and the distance along the X-axis represents the distance from the transducer. A-mode 
ultrasound allows very precise measurements of distance between reflective interfaces, thereby 
finding utility in ophthalmology. Depth resolution is optimal but lateral resolution is null and 
temporal resolution is absent. 
B-mode (brightness mode) ultrasonography is similarly archaic, but the principle of B-mode 
imaging is used in M-mode and 2D imaging. With B-mode imaging, the reflected sound waves are 
ascribed a pixel of a specific intensity along a vertical straight line. The brightness of a pixel, 
represented along a gray scale, is proportional to the amplitude of the echo, while its position on 
the screen corresponds to the depth from which the echo originated. Depth resolution is optimal but 
lateral resolution is null and temporal resolution is absent. 
M-mode (motion mode) ultrasonography (also known as TM-mode or Time-Motion mode) produces 
a tracing that represents the depth along the vertical axis, and time along the horizontal axis. It 
utilizes B-mode imaging technology along a single ultrasound beam, but acquires multiple linear B-
mode images at a very high frequency and displays them along the horizontal time axis. 
Consequently, pixels that represent specific tissue interfaces of moving structures appear to move 
closer or further away from the transducer, producing a wavy line. M-mode echocardiography has 
been extensively utilized in the past and is still used in veterinary practice, because the high axial 
resolution allows precise measurements of thickness, diameters and movement of cardiac 
structures. Because temporal resolution is also very high, exact measurements of the duration of 
cardiac events and their precise timing can be obtained. (New generation high-end ultrasound 
machines have processing capabilities that allow 2D imaging at frame rates that approach those 
obtainable in M-mode imaging - 150-200Hz - allowing similar axio-temporal resolution to M-mode 
imaging.) Depth resolution is optimal, lateral resolution is null and the temporal resolution is 
optimal. 
Real time (real-time) or 2D-mode (two-dimensional) ultrasonography refers to the ability to 
visualize the echocardiographic image in motion. The image that appears on the screen is 
composed of a series of B-mode lines, adjacent to each other, resulting in a gray scale rectangular 
or fan-shaped field of view. The simultaneous display of adjacent B-mode lines results in an 
anatomical cross-section of the organ being examined. Every B-mode line is obtained and 
processed virtually instantaneously, producing the image that appears on the screen. Every line of 
the image effectively represents one monodimensional ultrasound beam. Every line remains on the 
screen until it is replaced by its successor. The frequency with which the images are projected 
depends on the depth of origin of the echoes (which determines the pulse repetition frequency). 
Therefore, the deeper the image field, the longer the time necessary for the return of echoes to the 
transducer and therefore the lower the frame rate. Depth resolution is optimal, lateral resolution is 
optimal and temporal resolution is variable and dependent on the frame rate. 
During a complete echocardiographic examination, both the M-mode and two-dimensional techniques are 
routinely used as they provide different and complementary information. It should be emphasized that the 
ultrasonographer must not rely or depend on only one echocardiographic technique. Each imaging technique 
can inform the clinician, allowing him or her to analyze the case from different points of view. However, with 
advances in image acquisition and processing, frame rates (which previously limited temporal and spatial 
resolution in two-dimensional imaging and prevented accurate measurements) have increased phenomenally, 
allowing clinicians to obtain many measurements with two-dimensional images that could previously only be 
obtained with M-mode imaging. 
M-mode imaging is often done with two-dimensional guidance because the two-dimensional images offer an 
immediate and more intuitive interpretation of morphology and spatial orientation of the various cardiac 
structures. More recently, "anatomic M-mode" has been developed allowing