Ebook Essential echocardiography - A companion to braunwald’s heart disease: Part 1

ESSENTIAL
ECHOCARDIOGRAPHY
A Companion to Braunwald’s Heart Disease
Scott D. Solomon, MD

The Edward D. Frohlich Distinguished Chair
Professor of Medicine, Harvard Medical School Director
Director, Noninvasive Cardiology
Brigham and Women’s Hospital
Boston, Massachusetts

Justina C. Wu, MD, PhD
Assistant Professor of Medicine
Co-Director, Noninvasive Cardiology
Brigham and Women’s Hospital
Boston, Massachusetts

Linda D. Gillam, MD, MPH, FACC, FASE, FESC
Dorothy and Lloyd Huck Chair
Department of Cardiovascular Medicine
Medical Director, Cardiovascular Service Line
Morristown Medical Center/Atlantic Health System
Morristown, New Jersey
Professor of Medicine
Sidney Kimmel Medical College
Thomas Jefferson University
Philadelphia, Pennsylvania

Illustration Editor

Bernard E. Bulwer, MD, FASE
Noninvasive Cardiovascular Research
Cardiovascular Division
Brigham and Women’s Hospital
Boston, Massachusetts


1600 John F. Kennedy Blvd.
Ste 1800
Philadelphia, PA 19103-2899

Essential Echocardiography

ISBN: 978-0-323-39226-6

Copyright © 2019 by Elsevier, Inc. All rights reserved.
No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical,
including photocopying, recording, or any information storage and retrieval system, without permission in writing
from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing
Agency, can be found at our website: www.elsevier.com/permissions.
This book and the individual contributions contained in it are protected under copyright by the Publisher (other
than as may be noted herein).

Notices
Knowledge and best practice in this field are constantly changing. As new research and experience broaden our
understanding, changes in research methods, professional practices, or medical treatment may become necessary.
Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using
any information, methods, compounds, or experiments described herein. In using such information or methods
they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility.
With respect to any drug or pharmaceutical products identified, readers are advised to check the most current

information provided (i) on procedures featured or (ii) by the manufacturer of each product to be administered,
to verify the recommended dose or formula, the method and duration of administration, and contraindications.
It is the responsibility of practitioners, relying on their own experience and knowledge of their patients, to make
diagnoses, to determine dosages and the best treatment for each individual patient, and to take all appropriate
safety precautions.
To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise,
or from any use or operation of any methods, products, instructions, or ideas contained in the material herein.
Library of Congress Cataloging-in-Publication Data
Names: Solomon, Scott D., editor. | Wu, Justina C., editor. | Gillam, Linda
  D., editor.
Title: Essential echocardiography : a companion to Braunwald’s Heart disease
  / [edited by] Scott D. Solomon, Justina C. Wu, Linda D. Gillam ;
  illustration editor, Bernard E. Bulwer.
Other titles: Essential echocardiography (2019) | Complemented by
  (expression): Braunwald’s heart disease. 10th edition.
Description: Philadelphia, PA : Elsevier, [2019] | Complemented by:
  Braunwald’s heart disease / edited by Douglas L. Mann, Douglas P. Zipes,
  Peter Libby, Robert O. Bonow, Eugene Braunwald. 10th edition. 2015. |
  Includes bibliographical references and index.
Identifiers: LCCN 2017045233 | ISBN 9780323392266 (pbk. : alk. paper)
Subjects: | MESH: Echocardiography
Classification: LCC RC683.5.E5 | NLM WG 141.5.E2 | DDC 616.1/207543--dc23 LC record
available at />
Executive Content Strategist: Dolores Meloni
Senior Content Development Specialist: Rae Robertson
Publishing Services Manager: Catherine Jackson
Project Manager: Tara Delaney
Design Direction: Renee Duenow

Printed in China

Last digit is the print number: 9 8 7 6 5 4 3 2 1


To Caren, Will, Katie, and Dan
Scott D. Solomon
To Tsu-ming and Grace
Justina C. Wu
To John, Laura, and Jack
Linda D. Gillam
To all who made this possible, and to my parents, Albertha and Joseph
Bernard E. Bulwer


Contributors

Theodore Abraham, MD

Meyer Friedman Distinguished Professor of Medicine
Director, Echocardiography
University of California at San Francisco
San Francisco, California

Vikram Agarwal, MD, MPH

Noninvasive Cardiovascular Imaging Program
Department of Medicine (Cardiology) and Radiology
Brigham and Women’s Hospital
Boston, Massachusetts

Lillian Aldaia, MD

Department of Cardiovascular Medicine
Morristown Medical Center, Gagnon
Cardiovascular Institute
Morristown, New Jersey

M. Elizabeth Brickner, MD

Professor
Department of Internal Medicine
Division of Cardiology
UT Southwestern Medical Center
Dallas, Texas

Bernard E. Bulwer, MD, FASE

Noninvasive Cardiovascular Research
Cardiovascular Division
Brigham and Women’s Hospital
Boston, Massachusetts

Romain Capoulade, PhD

Patrycja Z. Galazka, MD

Division of Cardiovascular Medicine
Brigham and Women’s Hospital
Boston, Massachusetts

Linda D. Gillam, MD, MPH, FACC, FASE, FESC

Dorothy and Lloyd Huck Chair
Department of Cardiovascular Medicine
Medical Director, Cardiovascular Service Line
Morristown Medical Center/Atlantic Health System
Morristown, New Jersey
Professor of Medicine
Sidney Kimmel Medical College
Thomas Jefferson University
Philadelphia, Pennsylvania

Alexandra Goncalves, MD, MMSc, PhD

Cardiovascular Division
Brigham and Women’s Hospital
Boston, Massachusetts
Department of Physiology and Cardiothoracic Surgery
University of Porto Medical School
Porto, Portugal

John Gorcsan III, MD

Professor of Medicine
Director of Clinical Research
Washington University in St. Louis
St. Louis, Missouri

John D. Groarke, MBBCh, MSc, MPH

Echocardiography
Massachusetts General Hospital

Boston, Massachusetts

Brigham and Women’s Hospital Heart and Vascular Center;
Cardio-Oncology Program
Dana-Farber Cancer Institute/Brigham and Women’s Hospital
Boston, Massachusetts

Maja Cikes, MD, PhD

Deepak K. Gupta, MD

Sarah Cuddy, MBBCh

Rebecca T. Hahn, MD, FACC, FASE

Assistant Professor
Department for Cardiovascular Diseases
University of Zagreb School of Medicine
University Hospital Centre Zagreb
Zagreb, Croatia
Brigham and Women’s Hospital Heart and
Vascular Center
Boston, Massachusetts

Jan D’hooge, PhD

Assistant Professor of Medicine
Division of Cardiovascular Medicine
Vanderbilt Translational and Clinical Cardiovascular Research Center
Vanderbilt University Medical Center

Nashville, Tennessee
Director of Interventional Echocardiography
Center for Interventional and Vascular Therapy
Columbia University Medical Center
New York, New York

Professor
Department of Cardiovascular Sciences
University of Leuven
Leuven, Belgium

Sheila M. Hegde, MD

Rodney H. Falk, MD

Carolyn Y. Ho, MD

Division of Cardiovascular Medicine
Brigham and Women’s Hospital
Boston, Massachusetts

Cardiovascular Division
Brigham and Women’s Hospital
Boston, Massachusetts
Cardiovascular Division
Brigham and Women’s Hospital
Boston, Massachusetts

vii

viii

Contributors

Stephen J. Horgan, MB, BCh, PhD
Cardiology Fellow
Morristown Medical Center
Morristown, New Jersey

Judy Hung, MD

Associate Director
Echocardiography
Division of Cardiology
Massachusetts General Hospital
Boston, Massachusetts

Eric M. Isselbacher, MD, MSc

Director, Healthcare Transformation Lab
Co-Director, Thoracic Aortic Center
Massachusetts General Hospital
Associate Professor of Medicine
Harvard Medical School
Boston, Massachusetts

Kurt Jacobsen, RDCS

Lead Sonographer

Echocardiography Lab
Brigham & Women’s Hospital
Boston, Massachusetts

Konstantinos Koulogiannis, MD
Associate Director
Cardiovascular Core Lab
Department of Cardiovascular Medicine
Morristown Medical Center
Morristown, New Jersey

André La Gerche, MBBS, PhD

Laboratory Head
Department of Sports Cardiology
Baker Heart and Diabetes Institute
Cardiologist
St. Vincent’s Hospital
Melbourne, Victoria, Australia
Visiting Professor
Department of Cardiovascular Medicine
KU Leuven
Leuven, Brabant, Belgium

Jonathan R. Lindner, MD

M. Lowell Edwards Professor of Cardiology
Knight Cardiovascular Institute and Oregon National
Prime Research Center
Oregon Health & Science University

Portland, Oregon

Dai-Yin Lu, MD

Instructor
National Yang-Ming University School of Medicine
Taipei, Taiwan
Visiting Scientist
Department of Cardiology
The Johns Hopkins University School of Medicine
Baltimore, Maryland

Judy R. Mangion, MD, FACC, FAHA, FASE
Associate Director of Echocardiography
Department of Cardiovascular Medicine
Brigham and Women’s Hospital
Boston, Massachusetts

Warren J. Manning, MD

Section Chief, Non-invasive Cardiac Imaging & Testing
Cardiovascular Division
Beth Israel Deaconess Medical Center
Professor of Medicine and Radiology
Harvard Medical School
Boston, Massachusetts

Leo Marcoff, MD

Director of Interventional Echocardiography

Department of Cardiovascular Medicine
Morristown Medical Center
Morristown, New Jersey
Assistant Professor of Medicine
Sidney Kimmel Medical College
Thomas Jefferson University
Philadelphia, Pennsylvania

Thomas H. Marwick, MBBS, PhD, MPH
Director and Chief Executive, Professor
Baker Heart and Diabetes Institute
Melbourne, Victoria, Australia

Federico Moccetti, MD

Oregon Health & Science University
Portland, Oregon
Cardiovascular Division
University Hospital Basel
Basel, Switzerland

Monica Mukherjee, MD, MPH
Assistant Professor of Medicine
Department of Cardiology
Johns Hopkins University
Baltimore, Maryland

Denisa Muraru, MD, PhD

Department of Cardiac, Thoracic, and Vascular Sciences

University of Padua
Padua, Italy

Jagat Narula, MD, DM, PhD

Associate Dean for Global Affairs and Professor
Departments of Medicine and Cardiology
Icahn School of Medicine at Mount Sinai
New York, New York

Faraz Pathan, MBBS

Imaging Cardiovascular Fellow
Menzies Institute for Medical Research
Hobart, Tasmania, Australia

Elke Platz, MD, MS

Assistant Professor
Department of Emergency Medicine
Brigham and Women’s Hospital
Harvard Medical School
Boston, Massachusetts

Jose Rivero, MD, RDCS

Cardiovascular Department
Brigham and Women’s Hospital
Boston, Massachusetts

ix
Mário Santos, MD, PhD

Sara B. Seidelmann, MD, PhD
Cardiovascular Division
Brigham and Women’s Hospital
Boston, Massachusetts

Keri Shafer, MD

Adult Congenital Heart Disease Cardiologist
Brigham and Women’s Hospital
Instructor
Boston Children’s Hospital
Harvard Medical School
Boston, Massachusetts

Amil M. Shah, MD, MPH

Assistant Professor of Medicine
Harvard Medical School
Associate Physician
Division of Cardiovascular Medicine
Brigham and Women’s Hospital
Boston, Massachusetts

Douglas C. Shook, MD, FASE

Chief, Division of Cardiac Anesthesia

Department of Anesthesiology, Perioperative and Pain Medicine
Brigham and Women’s Hospital, Harvard Medical School
Boston, Massachusetts

Scott D. Solomon, MD

The Edward D. Frohlich Distinguished Chair
Professor of Medicine, Harvard Medical School
Director, Noninvasive Cardiology
Brigham and Women’s Hospital
Boston, Massachusetts

Jordan B. Strom, MD

Division of Cardiovascular Disease
Beth Israel Deaconess Medical Center
Instructor in Medicine
Harvard Medical School
Boston, Massachusetts

Timothy C. Tan, MBBS, PhD

Clinical Associate Professor
Department of Cardiology
Westmead Hospital
University of Sydney
Westmead, Australia
Conjoint Associate Professor
Department of Cardiology
Blacktown Hospital, Western Sydney University

Blacktown, Australia

Eliza P. Teo, MBBS

The Department of Cardiology
Royal Melbourne Hospital
Melbourne, Australia

Seth Uretsky, MD, FACC

Medical Director of Cardiovascular Imaging
Department of Cardiovascular Medicine
Morristown Medical Center
Morristown, New Jersey
Professor of Medicine
Sidney Kimmel School of Medicine
Thomas Jefferson University
Philadelphia, Pennsylvania

Rory B. Weiner, MD

Inpatient Medical Doctor
Cardiology Division
Massachusetts General Hospital;
Assistant Professor of Medicine
Harvard Medical School
Boston, Massachusetts

Leah Wright, BAppSc

Baker Heart and Diabetes Institute
Melbourne, Victoria, Australia

Justina C. Wu, MD, PhD

Assistant Professor of Medicine
Co-Director, Noninvasive Cardiology
Brigham and Women’s Hospital
Boston, Massachusetts

Contributors

Faculty of Medicine
Department of Physiology and
Cardiothoracic Surgery
Cardiovascular R&D Unit
University of Porto
Department of Cardiology
Porto Hospital Center
Porto, Portugal


Preface

Echocardiography, or cardiac ultrasound, is the most commonly used imaging technique to visualize the heart and great vessels. It remains an essential
tool for cardiovascular evaluation and management despite the emergence
of other imaging techniques such as cardiac magnetic resonance, computed
tomography, and nuclear imaging (SPECT and PET). Echocardiography
has proven diagnostic and prognostic value in the vast majority of cardiovascular diseases. Compared to other techniques, it is relatively noninvasive, inexpensive, and has none of the harmful effects of ionizing radiation.
Because it is increasingly portable and available in virtually any clinical setting, it may be used by a wide variety of practitioners, including cardiologists, intensivists, emergency physicians, anesthesiologists, and others.

The practice of echocardiography requires a strong knowledge of the
physical principles underlying ultrasound, an understanding of cardiac
anatomy and physiology, and an appreciation of the ultrasonic appearance
of both normal variants and different cardiovascular diseases. Moreover,
echocardiography, at its core, is a hands-on technique in which obtaining
high-quality images is dependent on the skill and training of the operator.
Essential Echocardiography: A Companion to Braunwald’s Heart Disease,
is designed as a textbook in echocardiography for anyone interested in
learning the technique, including practicing cardiologists, cardiology fellows, sonographers, anesthesiologists, critical care physicians, emergency

physicians, radiologists, residents, and medical students. The text is
designed to be simple enough to serve as an introduction to the field,
yet comprehensive enough to serve as a reference for experienced practitioners. Written by expert echocardiographers and sonographers with an
emphasis on the practical rather than the esoteric, the book focuses on
the basic principles of anatomy, physiology, and the hands-on approaches
necessary to acquire and interpret echocardiographic images with a rigorous focus on clinical care. The abundant illustrations, most of which are
also available on Expert Consult, underscore the importance of visual
learning in echocardiography. The images selected comprise an extensive
collection of classic and clear examples, representing decades of experience over multiple institutions and also recent advances in the field.
Echocardiography remains a vital and evolving technology. As a part
of the Heart Disease family, Essential Echocardiography will ensure that
students and practitioners of cardiology will have the tools and skills necessary to apply ultrasonic imaging to the care of cardiac patients.
Scott D. Solomon, MD
Justina C. Wu, MD, PhD
Linda D. Gillam, MD, MPH, FACC, FASE, FESC
Eugene Braunwald, MD

xi

Braunwald’s Heart Disease
Family of Books
BRAUNWALD’S HEART DISEASE COMPANIONS
BHATT
Cardiovascular
Intervention

DE LEMOS AND
OMLAND
Chronic Coronary
Artery Disease

ANTMAN AND
SABATINE
Cardiovascular
Therapeutics

ISSA, MILLER,
AND ZIPES
Clinical Arrhythmology
and Electrophysiology

xxiii


xxiv
BAKRIS AND
SORRENTINO
Hypertension

MCGUIRE AND
MARX
Diabetes in Cardiovascular
Disease

KORMOS AND
MILLER
Mechanical Circulatory
Support

MANN AND
FELKER
Heart Failure

MORROW
Myocardial Infarction

Braunwald’s Heart Disease Family of Books

BALLANTYNE
Clinical Lipidology


xxv

OTTO AND
BONOW
Valvular Heart Disease

BRAUNWALD’S HEART DISEASE REVIEW AND ASSESSMENT

LILLY
Braunwald’s Heart Disease

CREAGER, BECKMAN,
AND LOSCALZO
Vascular Medicine

Braunwald’s Heart Disease Family of Books

BLUMENTHAL, FOODY,
AND WONG
Preventative Cardiology


xxvi

Braunwald’s Heart Disease Family of Books

BRAUNWALD’S HEART DISEASE IMAGING COMPANIONS
TAYLOR

Atlas of Cardiovascular
Computer Tomography

KRAMER AND HUNDLEY
Atlas of Cardiovascular
Magnetic Resonance
Imaging

ISKANDRIAN AND

GARCIA

Atlas of Nuclear
Cardiology


section

I

Principles of Ultrasound
and Instrumentation

1

Physical Principles of Ultrasound
and Generation of Images
Maja Cikes, Jan D’hooge, Scott D. Solomon

INTRODUCTION
Ultrasound imaging is ubiquitous in medical practice and is used to image
all regions of the body, including soft tissues, blood vessels, and muscles.
The machines used for ultrasound imaging range from small hand-held
ultrasound devices no bigger than a smartphone to more elaborate and
complex systems capable of advanced imaging techniques such as threedimensional (3D) imaging. Although imaging of the heart and great vessels
has traditionally been referred to as “echocardiography,” the fundamental physical principles of image generation are common to all ultrasound
devices. These principles should be familiar to the end-user because they
are essential to understanding the utility and limitations of ultrasound and
to the interpretation of ultrasound images and can help optimize the use of
ultrasound systems to obtain the highest-quality images.

GENERATION OF IMAGES BY ULTRASOUND
The generation of images by ultrasound is based on the pulse-echo principle.1-3 It is initiated by an electric pulse that leads to the deformation of
a piezoelectric crystal housed in a transducer. This deformation results in
a high-frequency (>1,000,000 Hz) sound wave (ultrasound), which can
propagate through a tissue when the transducer is applied, resulting in an
acoustic compression wave that will propagate away from the crystal through
the soft tissue at a speed of approximately 1530 m/s. As with all sound
waves, each compression is succeeded by decompression: the rate of these
events defines the frequency of the wave. In diagnostic ultrasound imaging,
this applied frequency is generally between 2.5 and 10 MHz, which is far
beyond the level audible by humans, and is thus termed ultrasound.
The principal determinants of the ultrasound wave are: (1) wavelength
(λ), which represents the spatial distance between two compressions
(and is the primary determinant of axial resolution, as defined later), (2)
frequency (f ), which is inversely related to wavelength, and (3) velocity
of sound (c), which is a constant for any given medium (Fig. 1.1A and
B).These three wave characteristics have a set relationship as c = λf. An
increase in the frequency (i.e., shortening of the wavelength) implies less
deep penetration due to greater viscous effects leading to more attenuation. As the acoustic wave travels through tissue, changes in tissue properties, such as tissue density, will induce disruption of the propagating
wave, leading to partial reflection (specular reflections) and scatter (backscatter) of its energy (Fig. 1.2, Box 1.1).4 Typically, specular reflections
originate from interfaces of different types of tissue (such as blood pool
and myocardium or myocardium and pericardium), whereas backscatter
originates from within a tissue, such as myocardial walls. In both cases,
reflections propagate backwards to a piezoelectric crystal, again leading
to its deformation, which generates an electric signal. The amplitude of
this signal (termed the radiofrequency [RF] signal) is proportional to the
amount of deformation of the crystal (i.e., the amplitude of the reflected

wave). This signal is then amplified electronically, which can be modified by the “gain” settings of the system that will amplify both signal

and noise. In addition to defining the amplitude of the returning signal, the depth of the reflecting structure can be defined according to the
time interval from emitting to receiving a pulse, which equals the time
required for the ultrasound to travel from the transducer to the tissue and
back. The data on amplitude and depth of reflection are used to form scan
lines, and the overall image construction is based on repetitive operations
of the previously mentioned procedures of image (scan line) acquisition
and (post-) processing. During image acquisition, transducers emit ultrasound waves in pulses of a certain duration (pulse length), at a certain rate,
termed the pulse repetition frequency (PRF), which is one of the determinants of the temporal resolution of an echo image (obviously limited by
the duration of the pulse-echo measurement [i.e., its determinants]), as
elucidated further (see Fig. 1.1C).
The data obtained from scan lines can be visually represented as A- or
B-mode images (Fig. 1.3). The most fundamental modality of imaging
RF signals is A-mode, where A = amplitude, in which such signals are
imaged as amplitude spikes at a certain distance from the transducer; however, because visualization of the A-mode signals is relatively unattractive,
A-mode is not used as an image display option; further processing is used
to create a B-mode (B = brightness) image in which the amplitudes are
displayed by a gray scale (see Fig. 1.3). To achieve such gray scale encoding,
multiple points of the signal (i.e., pixels) are, based on the local amplitude
of the signal, designated with a number that further represents a color on
the gray scale. The B-mode dataset can then be displayed as an M-mode
(M = motion) image, which displays the imaged structures in one dimension over time (distance of the imaged structures from the transducer is
shown on the y-axis, and time is recorded on the x-axis; optimal for assessments requiring high temporal resolution and for linear measurements) or
as a 2D image. By convention, strong, high-amplitude reflections are given
a bright color and weak, low-amplitude reflections are dark (Box 1.2).
Another point in processing the RF signal overcomes a potential technical limitation of echocardiography; namely, reflections from tissues
more distant from the transducer are inherently smaller in amplitude, due
to attenuation (see Box 1.1). In practice, this implies that the segments
of the ultrasound image depicting, for example, the atria in the apical
views would be less bright than the myocardium. However, attenuation
correction can compensate for this effect, automatically amplifying the

signals from deeper segments, defined as automatic time-gain compensation (TGC) (Fig. 1.4). In addition to the automatic TGC, most systems
are equipped with TGC sliders that enable modification of the automated
TGC by the operator during image acquisition. Because the attenuation
effect can be variable among patients, the acquisition of echocardiographic
images should commence with a neutral setting of the sliders, which are
then individually modified according to the patient and the current echocardiographic view. Of note, attenuation cannot be corrected for after

1


2

Principles of Ultrasound and Instrumentation

I

A

B

C
FIG. 1.1  (A and B) Depiction of an (ultra)sound wave as a sine wave. The wave propagates through tissue at a given wavelength that is determined by frequency (to which it is inversely

related) and at a given amplitude that quantifies the amount of energy (i.e., the pressure change) transported by the wave. For sound waves, frequency is observed as pitch, whereas
amplitude is observed as loudness of the tone. (C) Pulse length (duration) is primarily determined by the transducer frequency, to which it is inversely related (e.g., higher-frequency transducers can emit pulses of shorter pulse length). These pulses are emitted at a certain rate, termed the pulse repetition frequency. (Courtesy of Bernard E. Bulwer, MD, FASE.)

image acquisition. The final step in image optimization, which can be
performed during post-processing, is log-compression—most often applied
in diagnostic imaging as the “dynamic range.” This method enables the
increase of image contrast by modifying the number of gray values, thus

leading to nearly black-and-white images (low dynamic range) or more
gray images (high dynamic range).2
Typically, the duration of the pulse-echo event is approximately
200 μs, taking into consideration the usual wave propagation distance during a cardiac examination (∼30-cm distance from the chest
wall to the roofs of the atria and back) and the speed of ultrasound
propagation through soft tissue. This implies that approximately 5000
pulse-echo measurements can be undertaken every second, while
approximately 180 of these measurements are performed in the construction of a typical 2D image of the heart, by emitting pulses in 180
different directions within a 90-degree scanning plane, reconstructing
one scan line for each transmitted pulse. In summary, a construction of one echocardiographic image requires approximately 36 ms
(180 measurements × 200 μs), which translates to approximately 28
frames created per second. However, the number of frames (i.e., the
frame rate) can be multiplied by various techniques, some of which are
implemented in most current systems, such as the multiline acquisition that constructs two or four lines in parallel, leading to a fourfold
increase in the 2D image frame rate.2 For more information on high
frame rate imaging, see Box 1.3.

Resolution of Echocardiographic Images
Resolution is defined as the shortest distance between two objects required
to discern them as separate. However, resolution in echocardiography,
being a dynamic technique, consists of two major components: spatial and
temporal resolution. Furthermore, spatial resolution mainly comprises axial
and lateral resolution, depending on the position of the objects relative to
the image line, and various determinants will influence each component of
image resolution (Figs. 1.5 to 1.7).1–3,5,6 Temporal resolution (i.e., frame
rate) represents the time between two subsequent measurements (i.e., the
ability of the system to discern temporal events as separate).
Axial resolution refers to resolution along the image line (i.e., two
objects located one behind another, relative to the image line) (see Fig.
1.6). Its principal determinant is pulse length (which is, similarly to

wavelength, inversely related to frequency), such that a shorter ultrasound pulse will allow for better axial resolution (typically 1.5 to 2
times the wavelength).2,6 Pulse length is predominantly defined by
the characteristics of the transducer: a higher-frequency transducer
provides shorter pulses, yielding better axial resolution. In practical
terms, a typical scanning frequency of 2.5 MHz implies a wavelength
of approximately 0.6 mm, at which an axial resolution of approximately 1 mm is obtained. However, higher frequencies have reduced
penetration due to more attenuation by soft tissue, implying that a
compromise between axial resolution and image depth needs to be


3
1

Physical Principles of Ultrasound and Generation of Images
FIG. 1.2  The interaction of the transmitted wave with an acoustic interface (i.e., cardiac structures). A segment of the transmitted wave is reflected at the interface,
while another part is transmitted through the tissue. Such a wave can be refracted, while the transmitted wave may also reflect and return to the transducer (thus carrying
information on signal amplitudes) as a specular reflection (mainly occurring at the interfaces of different types of tissue, such as myocardium and pericardium), or as backscatter
reflection (mainly originating from within the myocardial walls). LV, Left ventricle; PM, papillary muscle; PSAX, parasternal short-axis view; RV, right ventricle. (Modified from
Bulwer BE, Shernan SK, Thomas J. Physics of echocardiography. In: Savage RM, Aronson S, Shernan SK, eds. Comprehensive textbook of perioperative transesophageal echocardiography. Philadelphia: Lippincott, Williams & Wilkins; 2009:15.)

made. Therefore high-resolution imaging is predominantly limited to
pediatric echocardiography, where transducers up to 10 to 12 MHz
can be used for infants, as opposed to 2.5- to 3-MHz transducers
typically used in adult echocardiography.
Lateral resolution refers to the spatial resolution perpendicular to the
beam (i.e., two objects located next to each other, relative to the image
line) (see Fig. 1.7). It is predominantly determined by beam width,
which depends on depth and the size of the transducer footprint (Box
1.4). Lateral resolution will thus be increased with a narrower beam
(i.e., larger transducer footprint and/or shallower scanning depths).

Elevation resolution—resolution perpendicular to the image line—is
somewhat similar to lateral resolution. In this case the determinant is the
dimension of the beam in the elevation direction (i.e., orthogonal to the 2D
scan plane). Elevation resolution is more similar to lateral in newer systems
with 2D array transducer technology (compared with 1D transducers).

Temporal resolution, as mentioned previously, is predominantly
determined by PRF, which is limited by the determinants of the duration of the pulse-echo event—the wave propagation distance (the distance from the chest wall to the end of the scanning plane) and the
speed of ultrasound propagation through soft tissue (which is considered constant). Frame rate can be increased either by reducing the field
of view (a smaller sector requires the formation of fewer image lines,
allowing for a faster acquisition of a single frame) or by reducing the
number of lines per frame (line density), controlled by a “frame rate”
knob on the system. Reduced line density jeopardizes spatial resolution because it sets the image lines further apart. There is an intrinsic
trade-off between the image field of view, spatial resolution, and temporal resolution and should be kept in mind as a potential shortcoming of the technique (Box 1.5). For advice on image optimization, see
Box 1.6.


4

Principles of Ultrasound and Instrumentation

I

BOX 1.1  Attenuation, Reflection, and Refraction of
Ultrasound Waves
The attenuation of soft tissue is typically expressed in decibel
per cm per MHz (i.e., dB/cm per megahertz), given that the
attenuation is dependent on both frequency and propagation
distance of the wave. A typical value for attenuation in soft tissues is 0.5 dB/cm per megahertz, implying that for 20-cm propagation (e.g., from the probe to the mitral annulus and back
for an apical transducer position) of a wave generated by a

common adult cardiac ultrasound transducer (i.e., 2.5 MHz) the
amplitude of the acoustic wave has decreased by 25 dB, meaning that the wave received back at the probe surface will—at
best (i.e., assuming perfect reflection and optimal focusing)—
have only 5% of the amplitude of the transmitted wave. When
doubling the frequency to 5 MHz (i.e., pediatric probe) the
total attenuation doubles to 50 dB, implying that only approximately 0.3% of the transmitted amplitude returns from 20 cm
deep, which can become difficult to detect. Hence the proper
choice of transducer is required based on the depth at which
structures need to be visualized.
Reflection and refraction of sound waves occur at structures of differing acoustic impedance (i.e., mass density and/or
compressibility) that are large compared with the wavelength
(i.e., significantly > 0.5 mm for a 2.5 MHz wave). In this case the
behavior of acoustic waves is very similar to optic (i.e., light)
waves: part of the energy is transmitted into the second medium
under a slightly different angle (i.e., the refracted wave) while
part of the energy is reflected (i.e., the reflected wave). As a
simple example, you can think of what you see when holding
your hand under water: your arm appears to make an angle
at the water surface. The reason is light wave refraction at the
water surface, and the exact same phenomenon exists for ultrasound waves. One may thus think that the posterior wall would
appear distorted (cf. your arm under water) due to the wave
being refracted at the septal wall interfaces. Although this is
true, in practice these refraction effects are—luckily—most often
negligible.

Phased Array and Matrix Array Transducers
As opposed to mechanically rotating transducers used in earlier echocardiography systems, contemporary 2D imaging is based on electronic beam steering. This is achieved by an array of piezoelectric
crystals (typically up to 128 elements), while the time delay between
their excitation enables emission of the ultrasound wave in various
directions across the scan plane and the generation of multiple scan

lines (Fig. 1.8). The sum of signals received by individual elements
translates to the RF signal for a certain transmission, a process referred
to as beam forming (Box 1.7), which is crucial for acquiring high-quality images. Three-dimensional imaging relies on matrix array transducers, which are based on a 2D matrix of elements, thus enabling the
steering of the ultrasound beam in three dimensions. This allows for
both simultaneous multiplanar 2D imaging, as well as for volumetric
3D imaging.2

FIG. 1.3  Generation of images by ultrasound. After an ultrasound pulse is emitted

by the piezoelectric crystals located in the transducer (upper left), it travels through tissue, reflects from structures, and propagates backwards to the transducer. The received
signals undergo processing and are displayed according to their amplitudes and depth
of reflection (upper right). The fundamental A-mode display images the signals as amplitude spikes (upper right). On B-mode, these amplitude spikes are translated to a gray
scale, such that the least reflective tissues (e.g., blood pool) are visualized as black (upper
right). B-mode images can further be displayed as a two-dimensional cross-sectional
image (bottom left) or in M-mode, which visualizes the imaged structures in one dimension over time (bottom right). Note that reflections with the highest amplitudes originate from tissue interfaces such as the myocardium and pericardium or blood pool and
myocardium (upper and lower panels). IVS, Interventricular septum; LV, left ventricle;
PW, posterior wall. (Courtesy of Bernard E. Bulwer, MD, FASE; Modified from Solomon
SD, Wu J, Gillam L, Bulwer B. Echocardiography. In: Mann DL, Zipes DP, Libby P, Bonow
RO, Braunwald E, eds. Braunwald’s heart disease: a textbook of cardiovascular medicine.
10th ed. Philadelphia: Elsevier; 2015:180.)

BOX 1.2  Color Maps
The pixel values range from 0 to 255 (i.e., 28) for an 8-bit system, where 0 typically represents black, 255 represents white,
and the intermediate numbers correspond to hues of gray,
which can be extended to a spectrum of, for example, 65,536
(216) nuances of gray for the current systems with 16-bit resolution images. Furthermore, contemporary ultrasound systems
also offer a choice of color maps, in which case these values correspond to hues of, for example, bronze or purple. Although
gray-scale color maps are most often used, there is no scientific
rationale for this and some people prefer to use other color
schemes; this thus remains a matter of personal preference.

Second Harmonic Imaging
Current ultrasound systems are based on fundamental and harmonic
imaging. In fundamental imaging the transducer listens for the ultrasound of equal frequency to the emitted wave. However, at higher
amplitudes of the transmitted wave, wave distortion may occur during
propagation, causing harmonic frequencies (multiples of the transmitted frequency), which can be received by the transducer when properly
implemented (Fig. 1.9). Such second harmonic images have significantly
improved signal-to-noise ratio and in particular improved endocardial
border definition. However, this comes at the cost of poorer axial resolution (due to longer transmitted pulses), which may cause some structures,
such as heart valves, to appear thicker on harmonic imaging. The transition between fundamental and harmonic imaging is achieved by the
selection of transmit frequency: lower frequencies automatically enable

harmonic imaging, which is discernible by both the transmit and receive
frequency displayed on the screen (e.g., 1.7/3.4 MHz), whereas a single
displayed frequency implies fundamental imaging.1,2,5

PRINCIPLES OF DOPPLER IMAGING
Although imaging of the morphology of cardiac structures is increasingly
complemented by other modalities such as magnetic resonance imaging
(MRI) or computed tomography (CT) imaging, the diagnostic role of
echocardiographic imaging in the evaluation of valvular function and
noninvasive assessment of hemodynamics remains fairly unique. Such
assessments are based on the Doppler principle, which allows for the


5
1

reflecting structures, across various depths of the scan sector. (From Bulwer BE, Shernan SK. Optimizing two-dimensional echocardiographic imaging. In: Savage RM, Aronson S,
Shernan SK, eds. Comprehensive textbook of perioperative transesophageal echocardiography. Philadelphia: Lippincott, Williams & Wilkins; 2009:59.)

BOX 1.3  High Frame Rate Imaging
Multiple approaches have been proposed to increase frame
rate (i.e., time resolution) of the echocardiographic recordings. Most high-end commercially available systems reconstruct
2 to 4 image lines from each transmitted pulse, but 3D imaging systems reconstructing up to 64 lines for each transmit are
commercially available. Although this “parallel beam forming”
results in better time resolution of the images, it typically comes
at the cost of reduced spatial resolution and/or signal-to-noise
ratio of the images. Finding the optimal compromise between
these parameters is a major challenge for all vendors of ultrasound equipment. Alternative imaging techniques to speed
up the acquisition process but with potentially less effects
on spatial resolution and signal-to-noise ratio (e.g., multiline
transmit and diverging wave imaging) are being developed.
Two popular approaches that are currently being explored are
“multiline transmit” imaging and “diverging wave” imaging.
For the former a number of pulse-echo measurements are done
in multiple directions in parallel, a challenge being to avoid
crosstalk between the simultaneously transmitted pulses. In
the latter technique the whole field of view (or a large part
of it) is insonified by a very wide (i.e., defocused) ultrasound
beam, allowing to reconstruct the whole image with a very
small number of transmits (i.e., 1 to 5). In this way, frame rate is
increased tremendously (up to 1 to 5 kHz), the challenge being
to preserve spatial resolution and contrast of the images (i.e.,
image quality). Despite these remaining challenges, fast imaging approaches will undoubtedly enter clinical diagnostics in
the years to come.

calculations of blood velocities within the heart or in blood vessels.1–3,5,6
The Doppler effect states that the frequencies of transmitted and received
waves differ when the acoustic source moves towards or away from the

observer (due to wave compression or expansion, depending on the
direction of motion) (Fig. 1.10). For example, this is noticed as a higherpitched sound of the siren as the ambulance approaches the observer,

compared with it moving away. The Doppler effect can be applied to
measuring blood (and tissue) velocities, by measuring the difference
between the frequency of emitted and received ultrasound, which will be
reflected off moving red blood cells. Should the blood cells be moving in
the direction of the transducer, the reflected waves will be compressed and
the frequency of the received ultrasound will be higher compared with the
emitted ultrasound. Conversely, the frequency of the received ultrasound
will be lower with blood cells moving away from the transducer. This difference between the emitted and received frequency is termed the Doppler
shift or Doppler frequency, which is directly proportional to the velocity of
the reflecting structures (red blood cells, i.e., blood flow):


fd = 2 ft v (cos θ) /c

where fd is the Doppler frequency, ft is the original transmitted ultrasound
frequency, v is the magnitude of the velocity of blood flow, θ stands for
the angle between the ultrasound beam and the blood flow (i.e., the angle
of incidence/the angle of insonation), and c is the velocity of ultrasound
through soft tissue (1530 m/s). The main limitation of the Doppler
equation is the angle of incidence, such that its increase decreases the
calculated velocity: cos 0 degrees = 1, which implies that data acquisition
with the ultrasound beam parallel to the direction of blood flow would be
ideal; conversely, cos 90 degrees = 0, implying that motion orthogonal to
the ultrasound beam cannot be detected regardless of the velocity magnitude. Practically, an angle lower than 20 degrees is considered adequate for
acceptable measurements (of note, there is no possibility of velocity overestimation due to this phenomenon). To optimize alignment, Doppler
imaging can be used in conjunction with 2D imaging, which allows for
optimal placement of the Doppler cursor prior to Doppler data acquisition. Furthermore, should the angle of incidence be known, it can be

corrected for in the Doppler equation of the velocity estimate by means of
a feature available on many ultrasound systems, usually termed angle correction. However, this is acceptable for laminar flow conditions (typically
in vascular ultrasound, in particular of nonstenosed vessels), whereas the
exact direction of flow within the heart is, in fact, unknown. For this reason, it is not recommended to use angle correction in cardiac ultrasound
(or if applied, use with caution and awareness of the issue).

Physical Principles of Ultrasound and Generation of Images

FIG. 1.4  Attenuation correction settings. Optimal settings of time-gain compensation (TGC) can provide a uniform display of signal intensity for echoes from similarly


6

Principles of Ultrasound and Instrumentation

I

FIG. 1.5  Components of spatial resolution. Lateral resolution refers to the spatial resolution perpendicular to the beam, axial resolution refers to resolution along the
image line, and elevation resolution is also perpendicular to the image line; however, its determinant is the dimension of the beam in the elevation direction. (Modified from
Bulwer BE, Shernan SK. Optimizing two-dimensional echocardiographic imaging. In: Savage RM, Aronson S, Shernan SK, eds. Comprehensive textbook of perioperative transesophageal echocardiography. Philadelphia: Lippincott, Williams & Wilkins; 2009:54.)

Continuous Wave Doppler

Pulsed Wave Doppler

The Doppler modalities used in echocardiography are pulsed wave (PW)
and continuous wave (CW) Doppler (Fig. 1.11), as well as color flow
mapping (color flow Doppler). In CW, separate piezoelectric crystals
continuously emit and receive ultrasound waves, and the difference
between the frequencies of these waves (the Doppler shift) is calculated

continuously. In PW Doppler, ultrasound is emitted in pulses, as is the
case with standard image acquisition. According to the Doppler equation,
the Doppler shift is translated to velocity, which is then displayed over
a certain time frame (determined by the sweep speed of the image), and
is termed the spectrogram. As red blood cells travel at different velocities
within the ultrasound beam, various receive frequencies will be detected,
implying that a spectrum of Doppler shifts will be calculated and displayed on the spectrogram—thus termed spectral Doppler (Fig. 1.12).
In CW the spectrum is rather broad due to the large sample volume,
which accounts for a wide range of detected velocities, as opposed to
PW. Although ultrasound is well beyond the limits of human hearing,
the frequencies of the Doppler shift for typical blood velocities are actually within the audible range and can be heard during an examination:
a higher-pitched sound corresponds to higher velocities (larger Doppler
shift), whereas lower velocities generate a lower-pitched sound (smaller
Doppler shift). Furthermore, because the ultrasound waves are emitted
(and received) continuously in CW (i.e., the ultrasound system is not
“waiting” for the reflection and return of the emitted pulse), the location of the reflected sound cannot be determined and therefore no spatial information is available by CW. However, all frequency shifts (i.e.,
velocities) along the beam are measured, which allows for high-velocity
measurements by CW, typically used in the assessment of high velocities
(turbulence) across the aortic valve in patients with aortic stenosis or in
the approximation of pulmonary artery pressure from the velocity of the
tricuspid regurgitation jet. As is the case in 2D imaging the attenuation
effect also takes place in CW, as a consequence of which velocities from
deeper tissue contribute less to the displayed signal (Fig. 1.13). For advice
on CW Doppler optimization, see Box 1.8.

As opposed to CW, in PW Doppler ultrasound is emitted and received
in a similar manner to 2D imaging: individual pulses are emitted as brief,
intermittent bursts. After emitting such a pulse, the transducer “listens” to
returning signals only during a short, defined time interval following pulse
emission. This time interval corresponds to the time required for the pulse to

reach a certain depth and travel back to the transducer. The depth is defined
by the sample volume—in practical terms, a cursor that the operator places at
a certain depth along the transmitted beam, on the superimposed 2D image;
technically, this implies adjusting the timing between signal emission and
reception.7 Furthermore, the previously mentioned pulse-echo measurement is repeated along a specific line, at a specific repetition rate, termed the
PRF (i.e., the number of pulses transmitted from transducer per second).
Such pulses require time to reflect and travel back to the transducer; thus the
interval at which they are transmitted has to be long enough for the ultrasound system to be able to discern whether the reflected signal originates
from the given pulse or a later one. Based on this concept the velocity of
blood can be measured at a specific location in the heart by PW, thereby
providing spatial information on flows. Therefore PRF represents the sampling rate of the ultrasound machine: higher blood velocities imply higher
Doppler shift frequencies, requiring a higher sampling rate to detect the shift
(Box 1.9). Notably, PRF should not be mistaken for the frequency of the
ultrasound wave: in analogy to music, the PRF denotes the rate at which a
certain note is repeated, whereas the ultrasound wave frequency corresponds
to the pitch of a certain note.5 The PRF is a principal determinant of the
maximal Doppler shift (i.e., the maximal velocity within the sample volume
that the ultrasound system can accurately quantify). This maximal velocity
is also referred to as the Nyquist frequency (or the Nyquist limit) and is the
maximal velocity that can be accurately interrogated within a certain sample
volume. It is directly related to PRF, which is inversely related to the distance
between the transducer and sample volume. The Nyquist limit equals onehalf of the PRF. When imaging flows with velocities higher than double the
PRF value, sampling of the waveform is inaccurate, disabling the accurate
assessment of velocities, which can be detected by the appearance of aliasing


7
1

Physical Principles of Ultrasound and Generation of Images

FIG. 1.6  Features of axial resolution are based on pulse duration (spatial pulse, length), which is predominantly defined by the characteristics of the transducer
(i.e., its frequency). (A) The two reflectors (echo 1 and echo 2) are located apart enough to be resolved by the separately returning echo pulses. (B) The two reflectors (echo 1
and echo 2) are located too close, and the returning echo pulses will merge. (C) An increase in the transducer frequency from 3 to 7 MHz will shorten the spatial pulse length
(and pulse duration), thus permitting the returning echoes from these reflectors to be resolved. (Courtesy of Bernard E. Bulwer, MD, FASE.)


8

Principles of Ultrasound and Instrumentation

I

FIG. 1.7  Lateral resolution is predominantly determined by beam width, such that a narrower beam will allow for greater lateral resolution.

BOX 1.4  Beam Width

BOX 1.6  Image Optimization General Points

As a first approximation the beam width can be calculated as:
1.22.λ. d/D with “λ” the wavelength, “d” the focal depth, and
“D” the dimension of the transducer footprint. The ratio of d/D
is called the f-number of the transducer. From the previous equation, it is clear that transducer size directly impacts the spatial
resolution for a given depth. Unfortunately, for cardiac applications, transducer footprint needs to remain limited (and hence
the spatial resolution) due to the limited size of the acoustic window towards the heart (i.e., the intercostal space). Although, for
example, fetal cardiac imaging is possible with a cardiac ultrasound probe, image resolution will intrinsically be much better
when using a large, curved array as used in obstetrics.

•For optimal spatial resolution, use highest possible transducer frequency
•For optimal temporal resolution, use narrowest possible
sector and highest frame rate setting (i.e., lowest line

density)
•Optimize depth and focus according to imaged structure;
use minimal depth settings
•Optimize gain and dynamic range settings to obtain optimal image contrast: start with a black blood pool, increasing gain to a minimal amount that allows for definition of
the heart structures
•Time gain compensation should be used to homogenize
the image at various depths; start at a neutral position of
the sliders

BOX 1.5  The Trade-off Between Temporal and Spatial
Resolution
The trade-off between spatial resolution, temporal resolution,
signal-to-noise ratio and field of view of the echocardiographic
data is intrinsic and application dependent. Indeed, when measuring, for example, the dimensions of a given cardiac structure, time resolution may be less critical and system settings
could be adjusted to get the best possible spatial resolution
and signal-to-noise ratio at the cost of time resolution. On the
other hand, when making a functional analysis of the heart
(e.g., when applying speckle tracking), improved time resolution may be important and justify reducing the overall image
quality. It is thus important to realize that optimal acquisition
settings are application dependent.

in the generated image. Aliasing occurs due to the inability of the system to
accurately determine the velocity or direction of flow at velocities exceeding the Nyquist limit (Fig. 1.14). To avoid aliasing, a higher PRF should be
used, although a lower PRF will enable a better estimation of the blood flow
velocity—thus the lowest PRF possible without introducing aliasing should
be used. Depending on the machine, the PRF adjustment is referred to as
“scale,” “velocity range,” or “Nyquist velocity.”1–3,5,6 In addition, the baseline
of the spectrogram should be shifted upwards in case of flow away from the

transducer and downwards in case of flow towards the transducer, allowing

for higher velocities to be measured. Finally, a lower or higher PRF needs to
be applied depending on the depth of the measured flow: to “reach” flows at
greater depths (further from the transducer) and carry the information back
to the receiver, a lower PRF needs to be used, compared with flows closer to
the transducer. In practice, this is particularly obvious when measuring pulmonary vein flow in the apical views: a dedicated “low PRF” button on the
ultrasound system can be helpful to obtain an instantaneous shift in PRF and
improve signal quality. In analogy, higher velocities can be sampled without
aliasing at sample volume positions closer to the transducer. For advice on
PW Doppler optimization, see Box 1.10 and Fig. 1.15.

Color Flow Doppler
Color Doppler processing is based on PW Doppler imaging technology; however, in color flow Doppler the time shift between subsequent
measurements is determined at multiple sample volumes along multiple
scan lines. The calculated velocities are linked to a preset color scheme
by means of a specific color map (displayed on the ultrasound image,
Fig. 1.16), according to which the direction of flow and its velocity
amplitudes can be determined. By convention, flow away from the transducer is colored in blue, whereas flow towards the transducer is coded
in red. The color flow Doppler data are displayed superimposed on a
2D or M-mode image, allowing for visualization of flow patterns with


9
1

ing a fairly wide scan sector (center). During ultrasound transmission the time delays in activating the piezoelectric crystals induce the sweep of the scan line over the scan plane
(left). During reception, the reflected echo signals are out of phase when received by each crystal and need to be shifted in time (i.e., phased) prior to summation and further
processing (right). (Courtesy of Bernard E. Bulwer, MD, FASE; Modified from Solomon SD, Wu J, Gillam L, Bulwer B. Echocardiography. In: Mann DL, Zipes DP, Libby P, Bonow RO,
Braunwald E, eds. Braunwald’s heart disease: a textbook of cardiovascular medicine. 10th ed. Philadelphia: Elsevier; 2015:180.)

BOX 1.7  Beam Forming

Phased array transducers enable steering and focusing of the
ultrasound beam simply by adjusting the electrical excitations
of the individual transducer elements (see Fig. 1.7, left panel).
Similarly, during reception, the received signals coming from
individual transducer elements will be delayed in time to correct for the differing time of flight of a given echo to the individual transducer elements as a result of the differences in path
length to each of these elements (see Fig. 1.7, right panel). The
former is referred to as “transmit focusing,” whereas the latter is “receive focusing.” Interestingly, during receive focusing,
one can dynamically adjust the focus point as one knows a priori from which depth echo signals are arriving at a given time
point after transmission given the sound velocity is known. As
such, the time delays applied to the signals coming from the
different elements is adjusted dynamically in time to optimally
focus the ultrasound beam at all depths. Similarly, given that
focusing works better close the probe (see Box 1.4), some elements near the edge of the probe can be switched off when
(receive) focusing close to the probe to reduce the effective
transducer size, thereby making its ability to focus worse. The
advantage of this approach is that the beam width becomes
more uniform as a function of depth and thus so does the
lateral image resolution. These beam-forming modalities are
referred to as “dynamic receive focusing” and “dynamic apodization,” respectively, and are implemented on all cardiac
ultrasound systems.

additional information on the spatial location of the flow, nature of the
flow (turbulence, direction of flow), geometry of potential connections
between the heart chambers or great vessels, etc. Due to the same basic
principles of PW Doppler, color flow Doppler is also subject to aliasing, whereas a high variance of velocity in a particular pixel is mostly
displayed as shades of green, which is indicative of turbulent flow. Similar
to PW Doppler, the appearance of aliasing can be reduced by increasing
the PRF (however, PRF is coupled to velocity resolution) or by reducing
the transmission frequency (rarely performed). The generation of a color

flow Doppler image requires more “computing” time, and, to retain an
acceptable temporal resolution, it is suggested that the region of color
flow imaging (i.e., the color box) is kept to the minimal size required.1,2,5
For advice on color flow Doppler optimization, see Box 1.11.

Doppler Echocardiography in the Assessment of
Hemodynamics
Doppler echocardiography is predominantly used for the assessment of
velocities of blood flow within the heart and great vessels, which are determined by the driving pressure gradients between these structures (i.e.,
across heart valves). Analogously, the measured velocities of blood flow
across a certain valve can be used in the assessment of pressure gradients between the relevant chambers: based on conservation of energy, the
Bernoulli equation defines the relation between pressures and velocities for
fluids in chambers separated by an orifice:
P1 − P2 = 1 ρ V22 − V12
2
Connective Acceleration

+ ρ ∫1 dv ds
+
dt
Flow Acceleration
2

R v
Viscous Friction

where P1 and P2 represent the pressures, and V1 and V2 represent the
velocities proximal and distal to the orifice.1
In daily practice a simplified form of the Bernoulli equation may be
used, not taking into account flow acceleration and viscous friction:

P1 − P2 = 1/2ρ V22 − V12
Velocities proximal to the stenosis (i.e., orifice) are usually rather low
(when comparing with those distal to the stenosis) and may thus generally
be ignored, which further simplifies the equation:





P1 − P2 = 4V 2

Some of the most frequent applications of the Bernoulli equation in
the assessment of hemodynamics include the evaluation of the peak systolic gradient across the aortic valve in aortic stenosis: Doppler echocardiography can assess the peak velocity of antegrade blood flow across the
stenosed aortic valve, whereas applying the modified Bernoulli equation

Physical Principles of Ultrasound and Generation of Images

FIG. 1.8  The phased array transducer technology. Current echocardiography transducers steer the ultrasound beam (also termed sweep) across the scan plane, thus creat-


10

Principles of Ultrasound and Instrumentation

I

FIG. 1.9  Tissue harmonic imaging. Tissue harmonic imaging allows for improved image quality by using second-order harmonics in which specific frequencies of ultrasound
induce tissue vibrations at twice the frequency. Listening for such higher frequencies of returning ultrasound allows for dramatic improvement of the signal-to-noise ratio.
Second-harmonic imaging provides images with clearly ameliorated tissue definition and less affected by acoustic noise and artifacts (right). (Courtesy of Bernard E. Bulwer, MD,
FASE; From Solomon SD, Wu J, Gillam L, Bulwer B. Echocardiography. In: Mann DL, Zipes DP, Libby P, Bonow RO, Braunwald E, eds. Braunwald’s heart disease: a textbook of

cardiovascular medicine. 10th ed. Philadelphia: Elsevier; 2015:181.)

FIG. 1.10  The Doppler principle and Doppler frequency shift. Ultrasound emitted from the transducer reflects off moving red blood cells and returns to the transducer:
if reflected from red blood cells moving in the direction of the transducer, the echo returns at a higher frequency (shorter wavelength) than the emitted ultrasound pulse (upper
left); conversely, if blood cells are moving away from the transducer, a lower-frequency echo will be reflected back to the transducer (lower left). The difference between the
transmitted and the returning frequency equals the Doppler shift, which is used by Doppler echocardiography systems to calculate velocities of blood flow. These velocities are
graphically displayed by spectral Doppler as a time velocity spectrum (spectrogram), where a positive Doppler shift (implying flow toward the transducer) is depicted above the
baseline, and a negative Doppler shift (flow away from the transducer) is drawn below the baseline (right). In color flow Doppler the direction of flow can be detected according
to the color-coded velocities. (Courtesy of Bernard E. Bulwer, MD, FASE; Modified from Solomon SD, Wu J, Gillam L, Bulwer B: Echocardiography. In Mann DL, Zipes DP, Libby P,
Bonow RO, Braunwald E, eds. Braunwald’s Heart Disease: A Textbook of Cardiovascular Medicine. 10th ed. Philadelphia: Elsevier; 2015:182.)


11
1

Physical Principles of Ultrasound and Generation of Images
FIG. 1.11  Comparison of continuous wave (CW) Doppler and pulsed wave (PW) Doppler. (Courtesy of Bernard E. Bulwer, MD, FASE; Modified from Solomon SD, Wu
J, Gillam L, Bulwer B. Echocardiography. In: Mann DL, Zipes DP, Libby P, Bonow RO, Braunwald E, eds. Braunwald’s heart disease: a textbook of cardiovascular medicine. 10th
ed. Philadelphia: Elsevier; 2015:182.)

FIG. 1.12 The properties of spectral Doppler. The velocity of blood flow is

graphically displayed on the y-axis, and time is on the x-axis. Flow direction can also
be determined, depending on the relation of the spectrogram to the baseline: flow
toward the transducer is imaged above and flow away from the transducer is imaged
below the baseline. The signal intensity reflects the quantity of red blood cells that are
moving at a specific velocity range. In continuous wave the spectrum is rather broad
due to the wide range of velocities detected by the beam, as opposed to pulsed wave
(which is imaged here). A4C, Apical four-chamber.

FIG. 1.13 The depth attenuation effect seen on continuous wave (CW)
Doppler in aortic stenosis. With minimal gain settings, it can be appreciated that
the velocities from deeper tissues contribute less to the spectrogram: the Doppler
signal from the aortic root is attenuated and much weaker than that from the left
ventricular outflow tract (LVOT). With higher Doppler gain (second heart cycle), the
effect is less obvious. A5C, Apical five-chamber.


12

Principles of Ultrasound and Instrumentation

I

BOX 1.8  Continuous Wave Doppler Optimization
Points
•Optimize beam alignment with the direction of measured
velocity (direction of flow)
•Optimize gain to create a uniform Doppler profile free
of “blooming”: to prevent loss of data due to insufficient
gain, start with an overemphasized image, decreasing the
gain to a minimal required amount
•Optimize the “compress” control (assigns a certain shade
of color to varying amplitudes): extreme values can affect
the quality of the spectral analysis
•The “low velocities reject” button discards the signals of
lower amplitude, providing a cleaner image and more
precise measurements
•The “filter” reduces noise occurring from reflectors originating from the myocardium and other heart structures

BOX 1.9  High Pulse Repetition Frequency Pulsed Wave
Doppler
High PRF PW Doppler is also optional on some systems and can
be recognized by the occurrence of several sample volumes
along the Doppler beam. The measurement concept is based
on the fact that the PW Doppler system knows exactly when
to sample the echo signal (i.e., at the sample volume). As such,
a new pulse can already be transmitted (to a more proximal/
distal sample volume) before the echoes of the original transmit have been received without inducing artifacts. Thus the PRF
(and Nyquist limit) can be increased by emitting one (or more)
new pulses prior to receiving the signal of the first pulse from
the expected depth. However, such construction of the spectrogram implies that the exact location of the origin of the signal
along the Doppler beam cannot be known.

allows for the estimation of the peak instantaneous transvalvular gradient,
relevant for the assessment of aortic stenosis severity. Another frequently
used example refers to the assessment of peak systolic right ventricular
and pulmonary artery pressure: it is derived by adding the peak velocity
of the tricuspid regurgitation jet, which indicates the pressure gradient
between the right ventricle and right atrium in systole, to the right atrial
pressure estimate (which can also be determined by echocardiography,
according to the diameter and respiratory collapse of the inferior vena
cava). However, the Bernoulli equation can be used in all cases in which
a velocity gradient is present: valvular stenosis or regurgitation, as well
as abnormal connections (ventricular septal defect, etc.). Importantly, it
should be kept in mind that Doppler echocardiography enables the measurement of velocity, from which pressures and flows are inferred—the
absolute pressures in cardiac chambers can only be measured invasively.1
Another physical principle that is frequently used in the assessment
of hemodynamics is the continuity of flow equation, which states that the
same volume/flow passes through different cross sections of a tube (i.e.,

the heart), assuming no loss of fluid (i.e., no shunt). This equation is typically applied in the assessment of volumes/flows and valve areas: by multiplying the cross-sectional area (CSA) of the interrogated orifice by the
time velocity integral (TVI, i.e., the integration of blood velocity across
an orifice during one cardiac cycle) at the corresponding level, the magnitude of flow can be assessed (Fig. 1.17).
Furthermore, because a CSA of a diseased valve may be difficult to
measure, valve area can be calculated by estimating the flow proximal to
the valve and the TVI at the level of the valve. A frequently used example
includes the assessment of aortic valve CSA in aortic stenosis: according
to the continuity equation, the flow through the left ventricular outflow
tract (LVOT) equals that through the aortic valve, that is:


TVILVOT × AreaLVOT = TVIAV × AreaAV → AreaLVOT =
(TVIAV × AreaAV ) / TVILVOT

FIG. 1.14 The explanation of aliasing based the “wagon wheel” example,

stemming from the wagon wheel illusion seen in old western motion pictures (an
example from sampling theory): envision a rotating clock hand—in the top panel, it
rotates at one revolution per minute. If one would “sample” the clock 4 times per
minute (every 15 seconds) by shooting a picture, one could easily “capture” the
motion of the clock, could comprehend that the direction of rotation is clockwise,
and could perceive the rate of rotation. However, if the rotational speed were to be
increased to two revolutions per minute, maintaining the sampling rate, one would
“capture” only the hand at 12 o’clock and 6 o’clock, still being able to discern the
rate of rotation, but not the direction (middle panel). Ultimately, if the revolution
velocity increased to three revolutions per minute (in the same direction), retaining the same sampling rate, the perceived rate of rotation would be one revolution per minute while the perceived direction would be counterclockwise (bottom
panel). In analogy to pulsed wave Doppler, at a certain sampling rate of the system, increasing velocities of blood flow cannot be assessed adequately, neither for
their velocity, nor direction of blood flow. (From Solomon SD. Echocardiographic
instrumentation and principles of Doppler echocardiography. In: Solomon SD, ed.
Essential echocardiography—a practical handbook with DVD. Totowa, NJ, Humana

Press; 2007:12.)

BOX 1.10  Pulsed Wave Doppler Optimization Points
•Optimize beam alignment and gain, use the compress,
reject and filter settings as for CW Doppler
•Position the sample volume with particular caution: even
slight changes can affect the measurements significantly
(Fig. 1.15)
•Shift the baseline upwards or downwards to use the entire
display for either forward or backward flow (useful in
unidirectional flows)
•Optimize the PRF: use as high as possible to detect high
velocities, avoiding aliasing
•Use low PRF for flows distant from the transducer
•Use high PRF with caution if the origin of flow is relevant
CW, Continuous wave, PRF, pulse repetition frequency.

Obviously, such a calculation is prone to pitfalls that are mainly
due to erroneous measurement of the LVOT diameter, suboptimal
positioning of the PW Doppler sample volume in the LVOT, or
malacquisition of the peak velocities by CW Doppler across the aortic
valve.
An overview of hemodynamic data that can be derived from Doppler
echocardiography is given in Fig. 1.18. More detailed explanations of
specific measurements and entities will be given in further chapters of
the book.


13
1

volume should be positioned at the tips of the mitral valve leaflets, which would correspond to the mitral inflow pattern shown under the letter “E.” As can be observed, even
small deviations from this position can dramatically impact the pattern as well as the obtained measurements, thus rendering an inaccurate assessment of diastolic function. LA,
Left atrium; LV, left ventricle; MV, mitral valve; RA, right atrium; RV, right ventricle. (Modified from Appleton CP, Jensen JL, Hatle LK, Oh JK. Doppler evaluation of left and right
ventricular diastolic function: a technical guide for obtaining optimal flow velocity recordings. J Am Soc Echocardiogr. 1997;10(3):271-292, with permission.)

FIG. 1.16  Color flow Doppler imaging. Color flow Doppler is superimposed on the two-dimensional image. By convention, blood flow with mean velocities traveling

toward the transducer is encoded in red, and mean velocities moving away from the transducer are color-coded in blue. Similar to other forms of PW Doppler, high velocities and
turbulent flow are subject to aliasing, which is in color flow Doppler depicted as a multicolored mosaic pattern (typically green and yellow). The color-velocity scale illustrates incremental velocities in both directions from the baseline, such that higher velocities appear in increasingly lighter hues. A4C, Apical four-chamber; BA RT, blue away - red toward; LA,
left atrium; LV, left ventricle; RA, right atrium; RV, right ventricle. (Courtesy of Bernard E. Bulwer, MD, FASE; Modified from Solomon SD, Wu J, Gillam L, Bulwer B. Echocardiography. In: Mann DL, Zipes DP, Libby P, Bonow RO, Braunwald E, eds. Braunwald’s heart disease: a textbook of cardiovascular medicine. 10th ed. Philadelphia: Elsevier; 2015:183.)

Physical Principles of Ultrasound and Generation of Images

FIG. 1.15  The effect of sample volume position on the mitral inflow pattern. For the assessment of left ventricular diastolic function, the pulsed wave Doppler sample