Ebook Textbook of clinical echocardiography (5th edition): Part 1

Look for These Other Titles in the Otto Echocardiography Family

/
9
9

Echocardiography Review Guide: Companion to the Textbook of Clinical
Echocardiography, Second Edition
Catherine Otto, Rebecca Schwaegler, and Rosario Freeman

r
i
h

The Practice of Clinical Echocardiography, Fourth Edition
Catherine Otto

a
t
r/

Practical Echocardiography Series
Series Editor: Catherine Otto

e
s

Volumes Included in This Series:
Advanced Approaches in Echocardiography
Linda Gillam and Catherine Otto

/r u

Intraoperative Echocardiography
Donald Oxorn

.t c
a
k

Echocardiography in Heart Failure
Martin St. John Sutton and Susan Wiegers

/: /

Echocardiography in Congenital Heart Disease
Mark Lewin and Karen Stout

h

s
tt p


f

Fi

th

Edition

o

C a t h e r i n e M. O t t ,

m

TEXTBOOK of CLINICAL
/
ECHOCARDIOGRAPHY
9
d

h
a
t
r/

J. Ward Kennedy-Hamilton Endowed Chair in Cardiology

e
s

Professor of Medicine

University of Washington School of Medicine;

/r u

Director, Heart Valve Disease Clinic

Associate Director, Echocardiography Laboratory

.t c
a
k

University of Washington Medical Center
Seattle, Washington

h

s
tt p

/: /

ri 9


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


TEXTBOOK OF CLINICAL ECHOCARDIOGRAPHY
ISBN: 978-1-4557-2857-2
Copyright © 2013, 2009, 2004, 2000, 1995 by Saunders, an imprint of Elsevier Inc.

­

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.

r
i
h

This book and the individual contributions contained in it are protected under copyright by the Publisher (other
than as may be noted herein).

a
t
r/

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.

.t c
a
k

e
s

/r u

Library of Congress Cataloging-in-Publication Data
Otto, Catherine M.
Textbook of clinical echocardiography / Catherine M. Otto.—5th ed.
p. ; cm.—(Endocardiography)
Includes bibliographical references and index.
ISBN 978-1-4557-2857-2 (alk. paper)
I. Title. II. Series: Endocardiography.
[DNLM: 1. Echocardiography. 2. Heart Diseases—ultrasonography. WG 141.5.E2]
616.1′207543—dc23


s
tt p

/: /

h

Executive Content Strategist: Dolores Meloni
Senior Content Development Specialist: Joan Ryan
Publishing Services Manager: Deborah Vogel
Project Manager: Brandilyn Flagg
Designer: Lou Forgione

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

/
9
9

2012044607


PREFACE
Echocardiography is an integral part of clinical cardiology with important applications in diagnosis, clinical
management, and decision making for patients with
a wide range of cardiovascular diseases. In addition
to examinations performed in the echocardiography
laboratory, ultrasound imaging is now used in a variety
of other clinical settings, including the coronary care
unit, intensive care unit, operating room, emergency
department, catheterization laboratory, and electrophysiology laboratory, both for diagnosis and for monitoring the effects of therapeutic interventions. There
continues to be expansion of echocardiographic applications, given the detailed and precise anatomic and
physiologic information that can be obtained with this

technique at a relatively low cost and with minimal risk
to the patient.
This textbook on general clinical echocardiography
is intended to be read by individuals new to echocardiography and by those interested in updating their
knowledge in this area. The text is aimed primarily
at cardiology fellows on their basic echocardiography rotation but also will be of value to residents and
fellows in general internal medicine, radiology, anesthesiology, and emergency medicine, and to cardiac
sonography students. For physicians in practice, this
textbook provides a concise and practical update.
The Textbook of Clinical Echocardiography is structured
around a clinical approach to echocardiographic diagnosis. First, a framework of basic principles is provided with chapters on ultrasound physics, normal
tomographic transthoracic and transesophageal views,
intracardiac flow patterns, indications for echocardiography, and evaluation of left ventricular systolic
and diastolic function. A chapter on advanced echocardiographic modalities introduces the concepts of
3D echocardiography, myocardial mechanics, contrast
echocardiography, and intracardiac echocardiography. Clinical use of these modalities is integrated into
subsequent chapters as appropriate. This framework
of basic principles then is built upon in subsequent
chapters, organized by disease category (for example,
cardiomyopathy or valvular stenosis), corresponding to
the typical indications for echocardiography in clinical
practice.
In each chapter, basic principles for echocardiographic evaluation of that disease category are
reviewed, the echocardiographic approach and differential diagnosis are discussed in detail, limitations
and technical considerations are emphasized, and

h

s
tt p

/: /

.t c
a
k

/
9
9

alternate diagnostic approaches are delineated. Schematic diagrams are used to illustrate basic concepts;
echocardiographic images and Doppler data show
typical and unusual findings in patients with each
disease process. Transthoracic and transesophageal
images, Doppler data, and advanced imaging modalities are used throughout the text, reflecting their use in
clinical practice. Tables are used frequently to summarize studies validating quantitative echocardiographic
methods.
A special feature of this book that grew out of my
experience teaching fellows and sonographers is The
Echo Exam section at the end of the book. This section serves as a summary of the important concepts in
each chapter and provides examples of the quantitative calculations used in the day-to-day clinical practice of echocardiography. The information in The
Echo Exam is arranged as lists, tables, and figures for
clarity. My hope is that The Echo Exam will also serve
as a quick reference guide when a review is needed and
in daily practice in the echocardiography laboratory.
In the fifth edition, the text of all the chapters has
been revised to reflect recent advances in the field,
the suggested readings have been updated, and the
majority of the figures have been replaced with recent

examples that more clearly illustrate the disease process. The use of 3D and transesophageal imaging now
is explicitly integrated into each chapter. Additional
tables providing clinical-echocardiographic correlation have been added to several chapters. New artist
drawn illustrations provide a clearer understanding
of normal and abnormal cardiac anatomy. Updated
guidelines for the use of echocardiography and recommendations for image acquisition and analysis are
summarized in tables and illustrated in figures in each
chapter. The online and electronic versions of the
book are further enhanced by videos linked to the figures in each chapter.
A selected list of annotated references is included
at the end of each chapter. These references are suggestions for the individual who is interested in reading
more about a particular subject. Additional relevant
articles can be found in the suggested readings. Of
course, an online medical reference database is the
best way to obtain more recent publications and to
obtain a comprehensive list of all journal articles on
a specific topic.
For additional clinical examples, practical tips for
data acquisition, and self-assessment questions, the

a
t
r/

r
i
h

/r u

e
s

v


vi

Preface

Echocardiography Review Guide, by Otto, Schwaegler, and
Freeman (2nd edition, Elsevier/Saunders, 2011), parallels the information provided in this textbook and provides numerous multiple choice review questions with
detailed answers. A more advanced discussion of the
impact of echocardiographic data in clinical medicine
is available in a larger reference book, The Practice of
Clinical Echocardiography, 4th edition (Otto [ed], 2012),
also published by Elsevier/Saunders, with online cases,
video images, and interactive multiple choice questions
on the Expert Consult web site. Those seeking additional
expertise using echocardiography in specific clinical settings should consider the Otto Practical Echocardiography Series (Elsevier/Saunders, 2012) that includes
Advanced Approaches in Echocardiography (Gillam and Otto),
Intraoperative Echocardiography (Oxorn), Echocardiography in
Heart Failure (St John Sutton and Wiegers), and Echocardiography in Congenital Heart Disease (Lewin and Stout).
Each of these concise books provides practical clinical
approaches with numerous illustrations.
It should be emphasized that this textbook (or any
book) is only a starting point or frame of reference
for learning echocardiography. Appropriate training in echocardiography includes competency in the

h

s
tt p

/: /

.t c
a
k

acquisition and interpretation of echocardiographic
and Doppler data in real time. Additional training is
needed for performance of stress and transesophageal
examinations. Further, echocardiography continues to
evolve so that as new techniques become practical and
widely available, practitioners will need to update their
knowledge. Obviously, a textbook cannot replace the
experience gained in performing studies on patients
with a range of disease processes, and still photographs
or selected online videos do not replace the need for
acquisition and review of real-time data. Guidelines
for training in echocardiography, as referenced in
Chapter 5, serve as the standard for determining clinical competency in this technique. Although this textbook is not a substitute for appropriate training and
experience, I hope it will enhance the learning experience of those new to the field and provide a review for
those currently engaged in the acquisition and interpretation of echocardiography. Every patient deserves
a clinically appropriate and diagnostically accurate
echocardiographic examination; each of us needs to
continuously strive toward that goal.
Catherine M. Otto, MD

/r u

e
s

a
t
r/

r
i
h

/
9
9


ACKNOWLEDGMENTS
Many people have provided input to each edition of
the Textbook of Clinical Echocardiography, and the book is
immeasurably enhanced by their contributions—not
all can be individually thanked here but my gratitude
extends to all of you. My special thanks go to the cardiac sonographers at the University of Washington
for the outstanding quality of their echocardiographic
examinations and for our frequent discussions of the
details of image acquisition and the optimal echocardiography examination. Their skill in obtaining superb
images provides the basis of many of the figures in this
book. My thanks to Pamela Clark, RDCS; Sarah Curtis, RDCS; Caryn D’Jang, RDCS; Michelle Fujioka,
RDCS; Carol Kraft, RDCS; Yelena Kovolenko,

RDCS; Carol Kraft, RDCS; Chris McKenzie, RDCS;
Amy Owens, RDCS; Joanna Shephard, RDCS; Becky
Schwaegler, RDCS; Yu Wang, RDCS; and Todd
Zwink, RDCS.
My gratitude extends to my colleagues at the University of Washington who shared their expertise and
helped identify images for the book, including Rosario
Freeman, MD; Don Oxorn, MD; Eric Krieger, MD;
Steve Goldberg, MD; David Owens, MD; and Karen
Stout, MD. The University of Washington Cardiology Fellows also provided thoughtful (and sometimes
humbling) insights with particular recognition to Jason

s
tt p

/: /

.t c
a
k

/
9
9

Linefsky, MD, and Elisa Zaragoza-Macias, MD. In
addition, my gratitude includes my colleagues from
around the world who generously provided images,
including Marcia Barbosa, MD, and Maria P. Nunes,
MD, Belo Horizonte, Brazil; and Nozomi Watanabe,
MD, Kawasaki University, Okayama, Japan. Appreciation is also extended to those individuals who kindly

gave permission for reproduction of previously published figures. Joe Chovan and Starr Kaplan are to
be commended for their skills as medical illustrators
and for providing such clear and detailed anatomic
drawings.
My most sincere appreciation extends to the many
readers who provided suggestions for improvement with
particular thanks to Franz Wiesbauer and the participants in the 123 sonography community whose detailed
input that helped shape the 5th edition of this book.
Many thanks to my editor at Elsevier, Dolores Meloni,
for providing the support needed to write this edition,
and to Joan Ryan, Brandilyn Flagg, Michael Fioretti,
and the production team for all the detail-oriented hard
work that went into making this book and online videos
a reality.
Finally, my most appreciative thanks to my husband
and daughter for their unwavering support in every
aspect of life.

a
t
r/

r
i
h

/r u

e
s

h

vii


GLOSSARY

Abbreviations Used in Figures, Tables, and Equations

2D = two-dimensional
3D = three-dimensional
A-long = apical long-axis
A-mode= amplitude mode (amplitude versus depth)
A = late diastolic ventricular filling velocity with atrial
contraction
A′ = diastolic tissue Doppler velocity with atrial
contraction
A2C = apical two-chamber
A4C = apical four-chamber
AcT = acceleration time
Adur = transmitral A-velocity duration
adur = pulmonary vein a-velocity duration
AF = atrial fibrillation
AMVL = anterior mitral valve leaflet
ant = anterior
Ao = aortic or aorta
AR = aortic regurgitation
AS = aortic stenosis
ASD = atrial septal defect

ATVL = anterior tricuspid valve leaflet
AV = atrioventricular
AVA = aortic valve area
AVR = aortic valve replacement
BAV = bicuspid aortic valve
BP = blood pressure
BSA = body surface area

.t c
a
k

/: /

/
9
9

DT = deceleration time
dyne · s · cm-5 = units of resistance
D-TGA, complete transposition of the great arteries
E = early-diastolic peak velocity
E ′ = early-diastolic tissue Doppler velocity
ECG = electrocardiogram
echo = echocardiography
ED = end-diastole
EDD = end-diastolic dimension
EDV = end-diastolic volume
EF = ejection fraction
endo = endocardium

epi = epicardium
EPSS = E-point septal separation
ES = end-systole
ESD = end-systolic dimension
ESV = end-systolic volume
ETT = exercise treadmill test

a
t
r/

r
i
h

e
s

/r u

Δf = frequency shift
f = frequency
FL = false lumen
Fn = near field
Fo = resonance frequency
Fs = scattered frequency
FSV = forward stroke volume
FT = transmitted frequency

c = propagation velocity of sound in tissue

CAD = coronary artery disease
CPB = cardiopulmonary bypass
cath = cardiac catheterization
Cm = specific heat of tissue
cm/s = centimeters per second
cm = centimeters
CMR = cardiac magnetic resonance imaging
CO = cardiac output
cos = cosine
CS = coronary sinus
CSA = cross-sectional area
CT = computed tomography
CW = continuous-wave
Cx = circumflex coronary artery

HCM = hypertrophic cardiomyopathy
HPRF = high pulse repetition frequency
HR = heart rate
HV = hepatic vein
Hz = Hertz (cycles per second)

D = diameter
DA = descending aorta
dB = decibels
dP/dt = rate of change in pressure over time
dT/dt = rate of increase in temperature over time

l = length
LA = left atrium
LAA = left atrial appendage

LAD = left anterior descending coronary artery
LAE = left atrial enlargement
lat = lateral

s
tt p

h

I = intensity of ultrasound exposure
IAS = interatrial septum
ID = indicator dilution
inf = inferior
IV = intravenous
IVC = inferior vena cava
IVCT = isovolumic contraction time
IVRT = isovolumic relaxation time
kHz = kilohertz

xiii


xiv

Glossary

LCC = left coronary cusp
LMCA = left main coronary artery
LPA = left pulmonary artery
LSPV = left superior pulmonary vein

L-TGA = corrected transposition of the great
arteries
LV = left ventricle
LV-EDP = left ventricular end-diastolic pressure
LVH = left ventricular hypertrophy
LVID = left ventricular internal dimension
LVOT = left ventricular outflow tract
M-mode = motion display (depth versus time)
MAC = mitral annular calcification
MI = myocardial infarction
MR = mitral regurgitation
MS = mitral stenosis
MV = mitral valve
MVA = mitral valve area
MVL = mitral valve leaflet
MVR = mitral valve replacement
n = number of subjects
NBTE = nonbacterial thrombotic endocarditis
NCC = noncoronary cusp
ΔP = pressure gradient
P = pressure
PA = pulmonary artery
PAP = pulmonary artery pressure
PCI = percutaneous coronary intervention
PDA = patent ductus arteriosus or posterior
descending artery (depends on context)
PE = pericardial effusion
PEP = preejection period
PET = positron-emission tomography
PISA = proximal isovelocity surface area

PLAX = parasternal long-axis
PM = papillary muscle
PMVL = posterior mitral valve leaflet
post = posterior (or inferior-lateral) ventricular wall
PR = pulmonic regurgitation
PRF = pulse repetition frequency
PRFR = peak rapid filling rate
PS = pulmonic stenosis
PSAX = parasternal short-axis
PV = pulmonary vein
PVC = premature ventricular contraction
PVD = pulmonary vein diastolic velocity
PVR = pulmonary vascular resistance
PVD = pulmonary vein diastolic velocity
PWT = posterior wall thickness

s
tt p

.t c
a
k

/: /

h

Q = volume flow rate
Q p = pulmonic volume flow rate
Q s = systemic volume flow rate

r = correlation coefficient
R = ventricular radius
RFR = regurgitant instantaneous flow rate
RA = right atrium

RAE = right atrial enlargement
RAO = right anterior oblique
RAP = right atrial pressure
RCA = right coronary artery
RCC = right coronary cusp
R e = Reynolds number
RF = regurgitant fraction
RJ = regurgitant jet
R o = radius of microbubble
ROA = regurgitant orifice area
RPA = right pulmonary artery
RSPV = right superior pulmonary vein
RSV = regurgitant stroke volume
RV = right ventricle
RVE = right ventricular enlargement
RVH = right ventricular hypertrophy
RVol =regurgitant volume
RVOT = right ventricular outflow tract

a
t
r/

/
9

9

r
i
h

s = second
SAM = systolic anterior motion
SC = subcostal
SEE = standard error of the estimate
SPPA = spatial peak pulse average
SPTA = spatial peak temporal average
SSN = suprasternal notch
ST = septal thickness
STJ = sinotubular junction
STVL = septal tricuspid valve leaflet
SV = stroke volume or sample volume
(depends on context)
SVC = superior vena cava

e
s

/r u

T½ = pressure half-time
TD = thermodilution
TEE = transesophageal echocardiography
TGA = transposition of the great arteries
TGC = time-gain compensation

Th = wall thickness
TL = true lumen
TN = true negatives
TOF = tetralogy of Fallot
TP = true positives
TPV = time to peak velocity
TR = tricuspid regurgitation
TS = tricuspid stenosis
TSV = total stroke volume
TTE = transthoracic echocardiography
TV = tricuspid valve

v = velocity
V = volume or velocity (depends on context)
VAS = ventriculo-atrial septum
Veg = vegetation
Vmax = maximum velocity
VSD = ventricular septal defect
VTI = velocity-time integral
WPW = Wolff-Parkinson-White syndrome
Z = acoustic impedance


Glossary

Symbols

Greek Name

Used for

Variable

Unit

α

alpha

Frequency

Mass

g

γ

gamma

Viscosity

Grams. Example: LV
mass

Δ

delta

Difference

Pressure

mm Hg

θ

theta

Angle

λ

lambda

Wavelength

μ

mu

Micro-

Millimeters of mercury,
1 mm Hg = 1333.2
dyne/cm2, where dyne
measures force in
cm-mg-s2

π

pi

Mathematical constant
(approx. 3.14)

ρ

rho

Tissue density

σ

sigma

Wall stress

τ

tau

Time constant of
ventricular relaxation

Variable

Unit

Definition

Amplitude

dB

Decibels = a logarithmic
scale describing
the amplitude
(“loudness”) of the
sound wave

Angle

degrees

Degree = (π/180)rad.
Example: intercept
angle

Area

cm2

s
tt p

Frequency
(f)

Hz

kHz
MHz

h

Length

.t c
a
k

Square centimeters.
A 2D measurement
(e.g., end-systolic
area) or a calculated
value (e.g., continuity
equation valve area)

/: /

Hertz (cycles per
second)
Kilohertz = 1000 Hz
Megahertz =
1,000,000 Hz

cm

Centimeter (1/100 m)

mm

Millimeter (1/1000 m or
1/10 cm)

/
9
9

Resistance dyne · s · cm-5 Measure of vascular
resistance
Time

s
ms
μs

Ultrasound
intensity

W/cm2

Velocitytime
integral
(VTI)

e
s

/r u

Volume

mW/cm2

a
t
r/

Velocity (v)

UNITS OF MEASURE

Definition

m/s
cm/s
cm

cm3
mL
L

Volume
flow rate
(Q)

L/min
mL/s

Wall stress dyne/cm2
kdyn/cm2
kPa

Second
Millisecond (1/1000 s)
Microsecond

r
i
h

Where watt (W) =
joule per second and
joule = m2 · kg · s-2
(unit of energy)
Meters per second
Centimeters per second
Integral of the Doppler
velocity curve (cm/s)
over time (s), in units
of cm

Cubic centimeters
Milliliter, 1 mL = 1 cm3
Liter = 1000 mL
Rate of volume flow
across a valve or in
cardiac output
L/min = liters per minute

mL/s = milliliters per
second
Units of meridional or
circumferential wall
stress
Kilodynes per cm2
Kilopascals where
1 kPa = 10 kdyn/cm2

xv


KEY EQUATIONS

/
9
9

Ultrasound Physics
f = cycles/s = Hz
Frequency
λ = c / f = 1.54/f (MHz)
Wavelength
υ = c × Δf/ [2FT (cosθ)]
Doppler equation
ΔP = 4V 2
Bernoulli equation
LV Imaging
SV = EDV − ESV
Stroke volume

EF( % ) = (SV / EDV) × 100 %
Ejection fraction
σ = PR/2Th
Wall stress
Doppler Ventricular Function
SV = CSA × VTI
Stroke volume
dP/dt = 32 mm Hg / time from 1 to 3 m/s of MR CW jet(sec)
Rate of pressure rise
MPI = (IVRT + IVCT) / SEP
Myocardial performance index
Pulmonary Pressures and Resistance
PAPsystolic = 4(VTR )2 + RAP
Pulmonary systolic pressure
PAPsystolic = [4(VTR )2 + RAP] − Δ PRV − PA
PAP (when PS is present)
PAPmean = Mean Δ PRV − RA + RAP
Mean PA pressure
PAPdiastolic = 4(VPR )2 + RAP
Diastolic PA pressure
PVR ≅ 10(VTR )/VTIRVOT
Pulmonary vascular resistance
Aortic Stenosis
Δ Pmax = 4(Vmax )2
Maximum pressure gradient (integrate over ejection
period for mean gradient)
AVA(cm2 ) = [π(LVOTD / 2)2 × VTILVOT ] / VTIAS-Jet
Continuity equation valve area
AVA(cm2 ) = [π(LVOTD / 2)2 × VLVOT ] / VAS-Jet
Simplified continuity equation

Velocity ratio = VLVOT /VAS-Jet
Velocity ratio
Mitral Stenosis
MVADoppler = 220 / T½
Pressure half-time valve area
Aortic Regurgitation
TSV = SVLVOT = (CSALVOT × VTILVOT )
Total stroke volume
FSV = SVMA = (CSAMA × VTIMA )
Forward stroke volume
RVol = TSV − FSV
Regurgitant volume
ROA = RSV / VTIAR
Regurgitant orifice area
Mitral Regurgitation
TSV = SVMA = (CSAMA × VTIMA )
Total stroke volume
(or 2D or 3D LV stroke volume)
FSV = SVLVOT = (CSALVOT × VTILVOT )
Forward stroke volume
RVol = TSV − FSV
Regurgitant volume
ROA = RSV/VTIAR
Regurgitant orifice area
PISA method
RFR = 2πr2 × Valiasing
Regurgitant flow rate
ROAmax = RFR / VMR
Orifice area (maximum)
RV = ROA × VTIMR

Regurgitant volume
Aortic Dilation
Predicted sinus diameter
Children (<18 years): Predicted sinus dimension = 1.02 + (0.98 BSA)
Adults (18-40 years): Predicted sinus dimension = 0.97 + (1.12 BSA)
Adults (>40 years): Predicted sinus dimension = 1.92 + (0.74 BSA)
Ratio = Measured maximum diameter / Predicted maximum diameter
Pulmonary (Q p) to Systemic (Q s) Shunt Ratio
Q p : Q s = [CSAPA × VTIPA ] / [CSALVOT × VTILVOT ]

s
tt p

.t c
a
k

a
t
r/

r
i
h

/r u

e
s

/: /

h

xvii


1

Principles of Echocardiographic Image
Acquisition and Doppler Analysis

ULTRASOUND WAVES
ULTRASOUND TISSUE INTERACTION

DOPPLER ECHOCARDIOGRAPHY

Doppler Velocity Data
Doppler Equation
Spectral Analysis
Continuous-Wave Doppler Ultrasound
Pulsed Doppler Ultrasound
Doppler Velocity Instrument Controls
Doppler Velocity Data Artifacts
Color Doppler Flow Imaging
Principles
Color Doppler Instrument Controls
Color Doppler Imaging Artifacts
Tissue Doppler

Reflection
Scattering
Refraction
Attenuation

TRANSDUCERS

Piezoelectric Crystal
Types of Transducers
Beam Shape and Focusing
Resolution

ULTRASOUND IMAGING MODALITIES

M-Mode
Two-Dimensional Echocardiography
Image Production
Instrument Settings
Imaging Artifacts
Three-Dimensional Echocardiography
Echocardiographic Imaging Measurements

n understanding of the basic principles of
ultrasound imaging and Doppler echocardiography is essential both during data acquisition
and for correct interpretation of the ultrasound information. Although, at times, current instruments provide instantaneous images so clear and detailed that
it seems as if we can “see” the heart and blood flow
directly, in actuality, we always are looking at images
and flow data generated by complex analyses of ultrasound waves reflected and backscattered from the
patient’s body. Knowledge of the strengths, and more
importantly, the limitations, of this technique is critical

for correct clinical diagnosis and patient management.
On the one hand, echocardiography can be used for
decision making with a high degree of accuracy in a
variety of clinical settings. On the other hand, if an
ultrasound artifact is mistaken for an anatomic abnormality, a patient might undergo needless, expensive,
and potentially risky other diagnostic tests or therapeutic interventions.
In this chapter, a brief (and necessarily simplified)
overview of the basic principles of cardiac ultrasound
imaging and flow analysis is presented. The reader

A

BIOEFFECTS AND SAFETY
Bioeffects
Safety

SUGGESTED READING

is referred to the Suggested Reading at the end of
the chapter for more information on these subjects.
Because the details of image processing, artifact formation, and Doppler physics become more meaningful with experience, some readers may choose to
return to this chapter after reading other sections of
this book and after participating in some echocardiographic examinations.

ULTRASOUND WAVES
Sound waves are mechanical vibrations that induce
alternate refraction and compression of any physical
medium through which they pass (Fig. 1-1). Like other
waves, sound waves are described in terms of (Table 1-1):
  

  

n
n
n
n

 requency: cycles per second, or hertz (Hz)
F
Velocity of propagation
Wavelength: millimeters (mm)
Amplitude: decibels (dB)

Frequency ( f ) is the number of ultrasound waves in a
1-second interval. The units of measurement are hertz,
1


Chapter 1  |  Principles of Echocardiographic Image Acquisition and Doppler Analysis

abbreviated Hz, which simply means cycles per second.
A frequency of 1000 cycles/s is 1 kilohertz (KHz), and
1 million cycles/s is 1 megahertz (MHz). Humans can
hear sound waves with frequencies between 20 Hz and

20 kHz; frequencies higher than this range are termed
ultrasound. Diagnostic medical ultrasound typically uses
transducers with a frequency between 1 and 20 MHz.
The speed that a sound wave moves through the
body, called the velocity of propagation (c), is different for
Wavelength
λ

Propagation velocity (m/s)

each type of tissue. For example, the velocity of propagation in bone is much faster (about 3000 m/s) than
in lung tissue (about 700 m/s). However, the velocity
of propagation in soft tissues, including myocardium,
valves, blood vessels, and blood is relatively uniform,
averaging about 1540 m/s.
Wavelength is the distance from peak to peak of an
ultrasound wave. Wavelength can be calculated by
dividing the frequency ( f in Hz) by the propagation
velocity (c in m/s).
/
λ=c f



(1-1)

Since the propagation velocity in the heart is constant at 1540 m/s, the wavelength for any transducer
frequency can be calculated (Fig. 1-2) as:

Amplitude (dB)

2



//
/
λ (mm) = [1540m s f (Hz)] 1000 mm/m

or as:


/
λ (mm) = 1.54 f

For example, the wavelength emitted by a 5 MHz
transducer can be calculated as:
1s
cycles/s = Hz

/
/
λ = 1540 m s ÷ 5,000,000 cycle s = 0.000308 m
= 0.308 mm

Figure 1–1  Schematic diagram of an ultrasound wave.

TABLE 1-1  Ultrasound Waves
Definition

Examples

Clinical Implications

Frequency (f)

The number of cycles per
second in an ultrasound
wave:
f = cycles/s = Hz

Transducer frequencies
are measured in MHz
(1,000,000 cycles/s).
Doppler signal frequencies
are measured in KHz (1000
cycles/s).

Different transducer frequencies
are used for specific clinical
applications because the
transmitted frequency affects
ultrasound tissue penetration,
image resolution, and the
Doppler signal.

Velocity of
propagation (c)

The speed that ultrasound

travels through tissue

The average velocity of
ultrasound in soft tissue is
about 1540 m/s.

The velocity of propagation is
similar in different soft tissues
(blood, myocardium, liver, fat,
etc.) but is much lower in lung
and much higher in bone.

Wavelength (λ)

The distance between
ultrasound waves:
λ = c/f = 1.54/f (MHz)

Wavelength is shorter
with a higher-frequency
transducer and longer
with a lower-frequency
transducer.

Image resolution is greatest
(about 1 mm) with a
shorter wavelength (higher
frequency).
Depth of tissue penetration
is greatest with a longer

wavelength (lower frequency).

Amplitude (dB)

Height of the ultrasound
wave or “loudness”
measured in decibels
(dB)

A log scale is used for dB.
On the dB scale, 80 dB
represents a 10,000fold and 40 dB indicates
a 100-fold increase in
amplitude.

A very wide range of amplitudes
can be displayed using a grayscale display for both imaging
and spectral Doppler.


Principles of Echocardiographic Image Acquisition and Doppler Analysis  |  Chapter 1

Wavelength is important in diagnostic applications
for at least two reasons:
  

n
n

  

I mage resolution is no greater than 1 to 2 wavelengths (typically about 1 mm).
The depth of penetration of the ultrasound
wave into the body is directly related to wavelength; shorter wavelengths penetrate a shorter
distance than longer wavelengths.

Thus, there is an obvious tradeoff between image
resolution (shorter wavelength or higher frequency
preferable) and depth penetration (longer wavelength
or lower frequency preferable).
The acoustic pressure, or amplitude, of an ultrasound wave indicates the energy of the ultrasound
signal. Power is the amount of energy per unit time.
Intensity (I) is the amount of power per unit area:
Intensity (I) = power2



(1-2)

This relationship shows that if ultrasound power is
doubled, intensity is quadruped. Instead of using direct
measures of pressure energy, ultrasound amplitude is
described relative to a reference value using the decibel

scale. Decibels (dB) are familiar to all of us as the standard description of the loudness of a sound. Decibels
are logarithmic units based on a ratio of the measured
amplitude (A2) to a reference amplitude (A1) such that:
/
dB = 20 log (A2 A1 )

Thus, a ratio of 1000 to 1 is:

20 × log (1000) = 20 × 3 = 60 dB



a ratio of 100 to 1 is:
20 × log (100) = 20 × 2 = 40 dB



and a ratio of 2 to 1 is:
20 × log (2) = 20 × 0.3 = 6 dB



A simple rule to remember is that a 6-dB change
represents a doubling or halving of the signal amplitude or that a 40-dB change represents a 100 times difference in amplitude (Fig. 1-3). If acoustic intensity is
used instead of amplitude, the constant 10 replaces 20
in the equation so that a 3 dB changes represents doubling, and a 20 dB change indicates a 100-fold difference in amplitude. Either of these decibel scales may
30

0

Wavelength (mm)

.3
0.5

Wavelength (resolution)
Penetration

20

.62
1.0

10

Penetration (cm)

.2
.44

(1-3)

Figure 1–2  Transducer frequency versus
wavelength and penetration of the ultrasound signal in soft tissue. Wavelength
has been plotted inversely to show that resolution increases with increasing transducer
frequency while penetration decreases. The
specific wavelengths for transducer frequencies of 1, 2.5, 3.5, 5, and 7.5 MHz are shown.

.5

1.5
0

0

0 1 2.5 3.5

5

7.5

10

15

20

Transducer frequency (MHz)

Figure 1–3  Graph of the decibel scale.
The logarithmic relationship between the
decibel scale (horizontal axis) and the amplitude ratio (vertical axis) is seen. A doubling
or halving of the amplitude ratio corresponds
to a 6-dB change, and a 100-fold difference
in amplitude corresponds to a 20-dB change.

100,000

Amplitude ratio

10,000
1000
100
10
2

1
0
0

6

20

40

60
Decibels

80

100

3


4

Chapter 1  |  Principles of Echocardiographic Image Acquisition and Doppler Analysis

be used to refer to transmitted or received ultrasound
waves or to describe attenuation effects. The advantages of the decibel scale are that a very large range
can be compressed into a smaller number of values,
and that low-amplitude (weak) signals can be displayed
alongside very high-amplitude (strong) signals. In an
echocardiographic image, amplitudes typically range

from 1 to 120 dB. The decibel scale is the standard format both for echocardiographic image display and for
the Doppler spectral display, although other amplitude
scales may be an option.

ULTRASOUND TISSUE INTERACTION
Propagation of ultrasound waves in the body to generate ultrasound images and Doppler data depends on
a tissue property called acoustic impedance (Table 1-2).
Acoustic impedance (Z ) depends on tissue density (ρ)
and on the propagation velocity in that tissue (c):

Z = ρc



(1-4)

Although the velocity of propagation differs between
tissues, tissue density is the primary determinant of
acoustic impedance for diagnostic ultrasound. Lung tissue has a very low density compared to bone, which
has a very high density. Soft tissues, such as blood and
myocardium, have much smaller differences in tissue density and acoustic impedance. Acoustic impedance determines the transmission of ultrasound waves
through a tissue; differences in acoustic impedance result
in reflection of ultrasound waves at tissue boundaries.
The interaction of ultrasound waves with the organs
and tissues of the body can be described in terms of
(Fig. 1-4):
  

n
n
n
n

 eflection
R
Scattering
Refraction
Attenuation

TABLE 1-2  Ultrasound Tissue Interaction
Definition

Examples

Clinical Implications

Acoustic
impedance (Z)

A characteristic of each
tissue defined by
tissue density (r) and
propagation of velocity
(c) as:
Z=r×c

Lung has a low density and
slow propagation velocity,
whereas bone has a high
density and fast propagation
velocity. Soft tissues have
smaller differences in
tissue density and acoustic
impedance.

Ultrasound is reflected from
boundaries between
tissues with differences in
acoustic impedance (e.g.,
blood versus myocardium).

Reflection

Return of ultrasound signal
to the transducer from a
smooth tissue boundary

Reflection is used to generate
2D cardiac images.

Reflection is greatest when
the ultrasound beam is
perpendicular to the tissue
interface.

Scattering

Radiation of ultrasound in
multiple directions from
a small structure, such
as blood cells

The change in frequency
of signals scattered from
moving blood cells is the
basis of Doppler ultrasound.

The amplitude of scattered
signals is 100 to 1000
times less than reflected
signals.

Refraction

Deflection of ultrasound
waves from a straight
path because of
differences in acoustic
impedance

Refraction is used in transducer
design to focus the
ultrasound beam.

Refraction in tissues results in

double image artifacts.

Attenuation

Loss in signal strength
due to absorption of
ultrasound energy by
tissues

Attenuation is frequency
dependent with greater
attenuation (less penetration)
at higher frequencies.

A lower-frequency transducer
may be needed for apical
views or in larger patients
on transthoracic imaging.

Resolution

The smallest resolvable
distance between two
specular reflectors on an
ultrasound image

Resolution has three
dimensions: along the length
of the beam (axial), lateral
across the image (azimuthal)

and in the elevational plane.

Axial resolution is most precise
(as small as 1 mm), so
imaging measurements are
best made along the length
of the ultrasound beam.


Principles of Echocardiographic Image Acquisition and Doppler Analysis  |  Chapter 1

Reflection
The basis of ultrasound imaging is reflection of the
transmitted ultrasound signal from internal structures.
Ultrasound is reflected at tissue boundaries and interfaces, with the amount of ultrasound reflected dependent on the:
  

n
n
  

 ifference in acoustic impedance between the
D
two tissues
Angle of reflection

Smooth tissue boundaries with a lateral dimension
greater than the wavelength of the ultrasound beam
act as specular, or “mirrorlike,” reflectors. The amount
of ultrasound reflected is constant for a given interface,

although the amount received back at the transducer
varies with angle because (like light reflected from a
mirror) the angle of incidence and reflection is equal.
Thus, optimal return of reflected ultrasound occurs
at a perpendicular angle (90°). Remembering this fact
is crucial for obtaining diagnostic ultrasound images.
It also accounts for ultrasound “dropout” in a twodimensional (2D) or three-dimensional (3D) image
when too little or no reflected ultrasound reaches the
transducer resulting from a parallel alignment between
the ultrasound beam and tissue interface.

Scattering
Scattering of the ultrasound signal, instead of reflection, occurs with small structures, such as red blood
cells suspended in fluid, because the radius of the cell
(about 4 µm) is smaller than the wavelength of the
ultrasound signal. Unlike a reflected beam, scattered

Transducers

Scattering from
moving blood cells

Reflection

Refraction

ultrasound energy may be radiated in all directions.
Only a small amount of the scattered signal reaches
the receiving transducer, and the amplitude of a scattered signal is 100 to 1000 times (40-60 dB) less than
the amplitude of the returned signal from a specular

reflector. Scattering of ultrasound from moving blood
cells is the basis of Doppler echocardiography.
The extent of scattering depends on:
  






  

n
n
n
n

 article size (red blood cells)
P
Number of particles (hematocrit)
Ultrasound transducer frequency
Compressibility of blood cells and plasma

Although experimental studies show differences in
backscattering with changes in hematocrit, variation
over the clinical range has little effect on the Doppler signal. Similarly, the size of red blood cells and
the compressibility of blood cells and plasma do not
change significantly. Thus, the primary determinant
of scattering is transducer frequency.
Scattering also occurs within tissues, such as the

myocardium, from interference of backscattered signals from tissue interfaces smaller than the ultrasound
wavelength. Tissue scattering results in a pattern of
speckles; tissue motion can be measured by tracking
these speckles from frame to frame, as discussed in
Chapter 4.

Refraction
Ultrasound waves can be refracted—deflected from a
straight path—as they pass through a medium with
a different acoustic impedance. Refraction of an
ultrasound beam is analogous to refraction of light
waves as they pass through a curved glass lens (e.g.,
prescription eyeglasses). Refraction allows enhanced
image quality by using acoustic “lenses” to focus the
ultrasound beam. However, refraction also occurs in
unplanned ways during image formation, resulting in
ultrasound artifacts, most notably the “double-image”
artifact.

Attenuation
Specular
reflector

Attenuation

Figure 1–4  Diagram of the interaction between ultrasound and body
tissues. Doppler analysis is based on the scattering of ultrasound in all
directions from moving blood cells with a resulting change in frequency of
the ultrasound received at the transducer. 2D imaging is based on reflection
of ultrasound from tissue interfaces (specular reflectors). Attenuation limits

the depth of ultrasound penetration. Refraction, a change in direction of the
ultrasound wave, results in imaging artifacts.

Attenuation is the loss of signal strength as ultrasound
interacts with tissue. As ultrasound penetrates into the
body, signal strength is progressively attenuated because
of absorption of the ultrasound energy by conversion
to heat, as well as by reflection and scattering. The
degree of attenuation is related to several factors,
including the:
  






  

n
n
n
n

 ttenuation coefficient of the tissue
A
Transducer frequency
Distance from the transducer
Ultrasound intensity (or power)

The attenuation coefficient (α) for each tissue is
related the decrease in ultrasound intensity (measured

5


6

Chapter 1  |  Principles of Echocardiographic Image Acquisition and Doppler Analysis

in dB) from one point (I1) to a second point (I2) separated by a distance (l) as described by the equation:


I2 = I1 e − 2αl

(1-5)

The attenuation coefficient for air is very high (about
1000×) compared to soft tissue so that any air between
the transducer and heart results in substantial signal
attenuation. This is avoided on transthoracic examinations by use of a water-soluble gel to form an airless
contact between the transducer and the skin; on transesophageal echocardiography (TEE) examination, attenuation is avoided by maintaining close contact between
the transducer and esophageal wall. The air-filled lungs
are avoided by careful patient positioning and the use of
acoustic “windows” that allow access of the ultrasound
beam to the cardiac structures without intervening lung
tissue. Other intrathoracic air (e.g., pneumomediastinum, residual air after cardiac surgery) also results in
poor ultrasound tissue penetration because of attenuation, resulting in suboptimal image quality.
The power output of the transducer is directly related
to the overall degree of attenuation. However, an increase

in power output may cause thermal and mechanical
bioeffects as discussed in Bioeffects and Safety, p. 27.
Overall attenuation is frequency-dependent such
that lower ultrasound frequencies penetrate deeper
into the body than higher frequencies. The depth of
penetration for adequate imaging tends to be limited
to approximately 200 wavelengths. This translates
roughly into a penetration depth of 30 cm for a 1-MHz
transducer, 6 cm for a 5-MHz transducer, and 1.5 cm
for a 20-MHz transducer, although diagnostic images
at depths greater than these postulated limits can be
obtained with state-of-the-art equipment. Thus, attenuation, as much as resolution, dictates the need for a particular transducer frequency in a specific clinical setting.
For example, visualization of distal structures from the
apical approach in a large adult patient often requires
a low-frequency transducer. From a TEE approach, the
same structures can be imaged (at better resolution) with
a higher-frequency transducer. The effects of attenuation are minimized on displayed images by using different gain settings at each depth, an instrument control
called time-gain (or depth-gain) compensation.

TRANSDUCERS
Piezoelectric Crystal
Ultrasound transducers use a piezoelectric crystal both
to generate and to receive ultrasound waves (Fig. 1-5).
A piezoelectric crystal is a material (such as quartz or a
titanate ceramic) with the property that an applied electric current results in alignment of polarized particles
perpendicular to the face of the crystal with consequent
expansion of crystal size. When an alternating electric
current is applied, the crystal alternately compresses

and expands, generating an ultrasound wave. The frequency that a transducer emits depends on the nature

and thickness of the piezoelectric material.
Conversely, when an ultrasound wave strikes the
piezoelectric crystal, an electric current is generated.
Thus, the crystal can serve both as a “receiver” and
as a “transmitter.” Basically, the ultrasound transducer
transmits a brief burst of ultrasound and then switches
to the “receive mode” to await the reflected ultrasound
signals from the intracardiac acoustic interfaces. This
cycle is repeated temporally and spatially to generate
ultrasound images. Image formation is based on the
time delay between ultrasound transmission and return
of the reflected signal. Deeper structures have a longer
time of flight than shallower structures, with the exact
depth calculated based on the speed of sound in blood
and the time interval between the transmitted burst of
ultrasound and return of the reflected signal.
The burst, or pulse, of ultrasound generated by the
piezoelectric crystal is very brief, typically 1 to 6 µs,
because a short pulse length results in improved axial
(along the length of the beam) resolution. Damping
material is used to control the ring-down time of the
crystal and, hence, the pulse length. Pulse length also
is determined by frequency because a shorter time
is needed for the same number of cycles at higher
frequencies. The number of ultrasound pulses per
second is called the pulse repetition frequency, or PRF.
The total time interval from pulse to pulse is called
the cycle length, with the percent of the cycle length
used for ultrasound transmission called the duty factor.
Ultrasound imaging has a duty factor of about 1%

compared to 5% for pulsed Doppler and 100% for
continuous-wave (CW) Doppler. The duty factor is a
Transducer
Damping
material

Ultrasound
Pulse
Acoustic
lens

Pulse
length

Cable

λ
Piezoelectric
crystal

Impedance
matching

Figure 1–5  Schematic diagram of an ultrasound transducer. The
piezoelectric crystal both produces and receives ultrasound signals, with
the electric input-output transmitted to the instrument via the cable. Damping material allows a short pulse length (improved resolution). The shape
of the piezoelectric crystal, an acoustic lens, or electronic focusing (with a
phased-array transducer) are used to modify beam geometry. The material
of the transducer surface provides impedance matching with the skin. The
ultrasound pulse length for 2D imaging is short (1-6ms), typically consisting of two wavelengths (λ). “Ring down”—the decrease in frequency and

amplitude in the pulse—depends on damping and determines bandwidth
(the range of frequencies in the signal).


Principles of Echocardiographic Image Acquisition and Doppler Analysis  |  Chapter 1

key element in the patient’s total ultrasound exposure
as discussed in Bioeffects and Safety, p. 27.
The range of frequencies contained in the pulse is
described as its frequency bandwidth. A wider bandwidth
allows better axial resolution because of the ability
of the system to produce a narrow pulse. Transducer
bandwidth also affects the range of frequencies that
can be detected by the system with a wider bandwidth,

which allows better resolution of structures distant
from the transducer. The stated frequency of a transducer represents the center frequency of the pulse.

Types of Transducers
The simplest type of ultrasound transducer is based
on a single piezoelectric crystal (Table 1-3). Alternate

TABLE 1-3  Ultrasound Transducers
Definition

Examples

Clinical Implications

Type

Transducer characteristics
and configuration
Most cardiac transducers
use a phased array of
piezoelectric crystals.

Transthoracic (adult and pediatric)
Nonimaging CW Doppler
3D echocardiography
TEE
Intracardiac

Each transducer type is
optimized for a specific
clinical application.
More than one transducer may be
needed for a full examination.

Transmission
frequency

The central frequency
emitted by the
transducer

Transducer frequencies
vary from 2.5 MHz for
transthoracic echo to 20 MHz
for intravascular imaging.

Ahigher-frequency transducer
provides improved resolution
but less penetration.
Doppler signals are optimal at a
lower transducer frequency
than used for imaging.

Power output

The amount of ultrasound
energy emitted by the
transducer

An increase in transmitted
power increases the
amplitude of the reflected
ultrasound signals.

Excessive power output
may result in bioeffects
measured by the mechanical
and thermal indexes.

Bandwidth

The range of frequencies
in the ultrasound pulse

Bandwidth is determined by

transducer design.

A wider bandwidth allows
improved axial resolution for
structures distant from the
transducer.

Pulse (or burst)
length

The length of the
transmitted ultrasound
signal

A higher-frequency signal can
be transmitted in a shorter
pulse length compared to a
lower-frequency signal.

A shorter pulse length improves
axial resolution.

Pulse
repetition
frequency
(PRF)

The number of
transmission-receive
cycles per second

The PRF decreases as imaging
(or Doppler) depth increases
because of the time needed
for the signal to travel from
and to the transducer.

PRF affects image resolution and
frame rate (particularly with
color Doppler).

Duty factor

The percentage of time
that ultrasound is
transmitted

Ranges from about 1% for
imaging to 5% for pulsed
Doppler to 100% for CW
Doppler

A higher duty factor means more
tissue exposure to ultrasound.

Focal depth

Beam shape and
focusing are used to
optimize ultrasound

resolution at a specific
distance from the
transducer.

Structures close to the
transducer are best visualized
with a short focal depth,
distant structures with a long
focal depth.

The length and site of a
transducer’s focal zone
is primarily determined
by transducer design, but
adjustment during the exam
may be possible.

Aperture

The surface of the
transducer face
where ultrasound
is transmitted and
received

A small nonimaging CW
Doppler transducer allows
optimal positioning and
angulation of the ultrasound
beam.

A larger aperture allows a more
focused beam.
A smaller aperture allows
improved transducer
angulation on TTE imaging.

7


8

Chapter 1  |  Principles of Echocardiographic Image Acquisition and Doppler Analysis

pulsed transmission and reception periods allow
repeated sampling along a single line, with the sampling rate limited only by the time delay needed for
return of the reflected ultrasound wave from the depth
of interest. An example of using the transducer for
simple transmission-reception along a single line is an
A-mode (amplitude versus depth) or M-mode (depth
versus time) cardiac recording when a high sampling
rate is desirable.
Formation of more complex images uses an array
of ultrasound crystals arranged to provide a 2D tomographic or 3D volumetric data set of signals. Each
element in the transducer array can be controlled
electronically both to direct the ultrasound beam
across the region of interest and to focus the transmitted and received signals. Echocardiographic imaging
uses a sector scanning format with the ultrasound signal
originating from a single location (the narrow end of
the sector), resulting in a fanlike shape of the image.

Sector scanning is optimal for cardiac applications
because it allows a fast frame rate to show cardiac
motion and a small transducer size (aperture or “footprint”) to fit into the narrow acoustic windows used for
echocardiography. Three-dimensional imaging transducers are discussed in Chapter 4.
Most transducers can provide simultaneous imaging and Doppler analysis, for example, 2D-imaging
and a superimposed color Doppler display. Quantitative Doppler velocity data are recorded with the image
“frozen” or with only intermittent image updates, with
the ultrasound crystals used to optimize the Doppler
signal. Although CW Doppler signals can be obtained
using two elements of combined transducer, use of a
dedicated nonimaging transducer with two separate
crystals (with one crystal continuously transmitting
and the other continuously receiving the ultrasound
waves) is recommended when accurate high-velocity
recordings are needed. The final configuration of a
Figure 1–6  Schematic diagram of beam geometry for an unfocused (left) and focused
(right) transducer. The length of the near zone
and the divergence angle in the far field depend
on transducer frequency and aperture. The
focal zone of a focused transducer can be adjusted, but beam width still depends on depth.
Side lobes (and grating lobes with phased-array
transducers) occur with both focused and unfocused transducers and, like the central beam,
are 3D.

transducer depends on transducer frequency (higherfrequency transducers are smaller) and beam focusing,
as well as the intended clinical use, for example, transthoracic versus TEE imaging.

Beam Shape and Focusing
An unfocused ultrasound beam is shaped like the light
from a flashlight, with a tubular beam for a short distance that then diverges into a broad cone of light (Fig.

1-6). Even with current focused transducers, ultrasound
beams have a 3D shape that affects measurement accuracy and contributes to imaging artifacts. Beam shape
and size depend on several factors, including:
  






  

n
n
n
n

 ransducer frequency
T
Distance from the transducer
Aperture size and shape
Beam focusing

Aperture size and shape and beam focusing can be
manipulated in the design of the transducer, but the
effects of frequency and depth are inherent to ultrasound physics. For an unfocused beam, the initial segment of the beam is columnar in shape (near field Fn)
with a length dependent on the diameter D of the
transducer face and wavelength (λ):
/
Fn = D2 4λ

(1-6)

For a 3.5-MHz transducer with a 5-mm diameter
aperture, this corresponds to a columnar length of 1.4
cm. Beyond this region, the ultrasound beam diverges
(far field), with the angle of divergence θ determined as:
/
sin θ = 1.22λ D



(1-7)

This equation indicates a divergence angle of 6°
beyond the near field, resulting in an ultrasound beam
width of about 4.4 cm at a depth of 20 cm for this 3.5MHz transducer. With a 10-mm diameter aperture, Fn

Unfocused
transducer

Focused
transducer

Near zone
Side
lobes

Beam
width
Focal
zone

Divergence
angle


Principles of Echocardiographic Image Acquisition and Doppler Analysis  |  Chapter 1

20
20
10 mm aperture
10
10
5 mm aperture

Divergence angle (degrees)

Length of near zone (cm)

30

Figure 1–7  Transducer frequency versus
near zone length and divergence angle.
Transducer frequency is shown on the horizontal axis with the length of the near zone shown
in yellow and divergence angle in blue for unfocused 5- (squares) and 10-mm (triangles)
diameter aperture transducers. Equations (1-6)
and (1-7) were used to generate these curves.

0

0
0

5

10

Transducer frequency (MHz)

Main
lobe

Side
lobe 1

Side
lobe 2

F
θ2
θ1

λ
P1

P2

would be 5.7 cm and beam width at 20 cm would be
about 2.5 cm (Fig. 1-7).
The shape and focal depth (narrowest point) of the
primary beam can be altered by making the surface of
the piezoelectric crystal concave or by the addition of
an acoustic lens. This allows generation of a beam with
optimal characteristics at the depth of most cardiac
structures, but again, divergence of the beam beyond
the focal zone occurs. Some transducers allow manipulation of the focal zone during the examination. Even
with focusing, the ultrasound beam generated by each
transducer has a lateral and an elevational dimension
that depends on the transducer aperture, frequency,
and focusing. Beam geometry for phased-array transducers also depends on the size, spacing, and arrangement of the piezoelectric crystals in the array.
In addition to the main ultrasound beam, dispersion
of ultrasound energy laterally from a single-crystal
transducer results in formation of side lobes at an angle
θ from the central beam where sin θ = m λ /D, and
m is an integer describing sequential side lobes (i.e., 1,
2, 3, and so on) (Fig. 1-8). Reflected or backscattered
signals from these side lobes may be received by the
transducer, resulting in image or flow artifacts. With
phased-array transducers, additional accessory beams
at an even greater angle from the primary beam,

Intensity

Main lobe

Side
lobe 1

θ2

θ1

0
Angular position

θ1

Side
lobe 2

θ2

Figure 1–8  Transducer beam side lobes. Top: This diagram shows that
side lobes occur at the points where the distances traversed by the ultrasound
pulse from each edge of the crystal face differ by exactly one wavelength. The
distance from the left edge of the crystal (P1) to the position of side lobe 1 is
exactly one wavelength (λ) longer than the distance from the extreme right
edge of the crystal (P2) to the position of side lobe 1. Bottom: The beam intensity plot formed by sweeping along an arc at focal length F.  (From Geiser EA:
Echocardiography: physics and instrumentation. In Skorton DJ, Schelbert AR,
Wolf GL, Brundage BH [eds]: Marcus Cardiac Imaging, 2nd ed. Philadelphia:
WB Saunders, 1996, p 280. Used with permission.)

9


10

Chapter 1  |  Principles of Echocardiographic Image Acquisition and Doppler Analysis

Acoustic
lens

Lateral
Slice thickness
(elevational)
Axial

A

Resolution components
in 3D space

B

Elevational profile of
ultrasound beam with depth

Figure 1–9  Axial, lateral, and elevational slice thickness in three dimensions for a phased-array transducer ultrasound beam. A, Axial resolution along
the direction of the beam is independent of depth; lateral resolution and elevational resolution are strongly depth dependent. Lateral resolution is determined by
transmit and receive focus electronics; elevational resolution is determined by the height of the transducer elements. At the focal distance, axial is better than
lateral and is better than elevational resolution. B, Elevational resolution profile with an acoustic lens across the transducer array produces a focal zone in the
slice thickness direction.  (From Bushberg JT, et al: The Essential Physics of Medical Imaging. Philadelphia: Lippincott Williams & Wilkins, 2002, Fig. 16-21).

termed grating lobes, also occur as a result of constructive interference of ultrasound wave fronts. Both the
side lobes and the grating lobes affect the lateral and
elevational resolution of the transducer.

Resolution
Image resolution occurs for each of three dimensions
(Fig. 1-9):
  

n
n
n
  

 xial resolution along the length of the ultrasound
A
beam
Lateral resolution side to side across the 2D image
Elevational resolution or thickness of the tomographic slice

Of these three, axial resolution is most precise, so
quantitative measurements are made most reliably using
data derived from a perpendicular alignment between
the ultrasound beam and structure of interest. Axial
resolution depends on the transducer frequency, bandwidth, and pulse length but is independent of depth
(Table 1-4). Determination of the smallest resolvable
distance between two specular reflectors with ultrasound
is complex but is typically about twice the transmitted
wavelength; higher-frequency (shorter-wavelength)
transducers have greater axial resolution. For example,
with a 3.5 MHz transducer, axial resolution is about
1 mm, versus 0.5 mm with a 7.5 MHz transducer. A
wider bandwidth also improves resolution by allowing a shorter pulse, thus avoiding overlap between the
reflected ultrasound signals from two adjacent reflectors.

Lateral resolution varies with the depth of the specular reflector from the transducer, primarily related to

TABLE 1-4 Determinants of Resolution
in Ultrasound Imaging
Axial Resolution

Transducer frequency
Transducer bandwidth
Pulse length
Lateral Resolution

Transducer frequency
Beam width (focusing) at each depth*
Aperture (width) of transducer
Bandwidth
Side and grating lobe levels
Elevational Resolution

Transducer frequency
Beam width in elevational plane
*Most important.

beam width at each depth. In the focal region where
beam width is narrow, lateral resolution may approach
axial resolution, and a point target will appear as a
point on the 2D image. At greater depths, beam width
diverges so a point target results in a reflected signal
as wide as the width of the beam, which accounts for
“blurring” of images in the far field. If the 2D image
is examined carefully, progressive widening of the

echo signals from similar targets along the length of


Principles of Echocardiographic Image Acquisition and Doppler Analysis  |  Chapter 1

ULTRASOUND IMAGING MODALITIES
M-Mode

Figure 1–10  Beam width effect on 2D imaging. 2D echocardiographic view of the LV from an apical approach. The effect of beam width
can be appreciated by comparing the length of reflections from point targets
near and at greater distances from the transducer as shown by the arrows.

the ultrasound beam can be appreciated (Fig. 1-10).
Erron­eous interpretations occur when the effects of
beam width are not recognized. For example, beam
width artifact from a strong specular reflector may
appear to be an abnormal linear structure. Other
factors that affect lateral resolution are transducer
frequency, aperture, bandwidth, and side and grating
lobe levels.
Resolution in the elevational plane is more difficult
to recognize on the 2D image but is equally important in clinical diagnosis. The thickness of the tomographic plane varies across the 2D image, depending
on transducer design and focusing, both of which
affect beam width in the elevational plane at each
depth. In general, cardiac ultrasound images have a
“thickness” of approximately 3 to 10 mm depending on depth and the specific transducer used. The
tomographic image generated by the instrument, in
effect, includes reflected and backscattered signals
from this entire thickness. Strong reflectors adjacent
to the image plane may appear to be “in” the image

plane because of elevational beam width. Even more
distant strong reflectors may appear superimposed on
the tomographic plane because of side lobes in the
elevational plane. For example, a linear echo in the
aortic lumen from an adjacent calcified atheroma may
look like a dissection flap. These principles of ultrasound imaging also apply to 3D echocardiography
(see Chapter 4).

Historically, cardiac ultrasound began with a singlecrystal transducer display of the amplitude (A) of
reflected ultrasound versus depth on an oscilloscope
screen. This A-mode display may still be shown on
the 2D image screen to aid the examiner in optimal
adjustment of the instrument controls. Repeated pulse
transmission-and-receive cycles allow rapid updating
of the amplitude-versus-depth information so that
rapidly moving structures, such as the aortic or mitral
valve leaflets, can be identified by their characteristic
timing and pattern of motion (Fig. 1-11).
With the time dimension shown explicitly on the
horizontal axis and each amplitude signal along the
length of the ultrasound beam converted to a corresponding gray-scale level, a motion (M) mode display is
produced. M-mode data are shown on the video monitor either “scrolling” or “sweeping” across the screen
at 50 to 100 mm/s. Two-dimensional (2D) imaging
allows guidance of the M-mode beam to ensure an
appropriate angle between the M line and the structures of interest.
Because only a single “line of sight” is included in
an M-mode tracing, the pulse repetition frequency
(PRF) of the transmission-and-receive cycle is limited only by the time needed for the ultrasound beam
to travel to the maximum depth of interest and back
to the transducer. Even a depth of 20 cm requires

only 0.26 ms (given a speed of propagation of 1540
m/s), allowing a PRF up to 3850 times per second.
In actual practice, sampling rates of about 1800
times per second are used. This extremely high sampling rate is valuable for accurate evaluation of rapid
normal intracardiac motion such as valve opening
and closing. In addition, continuously moving structures, such as the ventricular endocardium, may be
identified more accurately when motion versus time,
as well as depth, is displayed clearly on the M-mode
recording. Other examples of rapid intracardiac
motion best demonstrated with M-mode imaging
include the high-frequency fluttering of the anterior
mitral leaflet in patients with aortic regurgitation and
the rapid oscillating motion of valvular vegetations.

Two-Dimensional Echocardiography
Image Production
A 2D echocardiographic image is generated from the
data obtained by electronically “sweeping” the ultrasound beam across the tomographic plane. For each
scan line, short pulses (or bursts of ultrasound) are
emitted at a PRF determined by the time needed for
ultrasound to travel to and from the maximum image
depth. The pulse repetition period is the total time

11


Chapter 1  |  Principles of Echocardiographic Image Acquisition and Doppler Analysis

3D/2D

M-mode

Distance

Time

Ao

A-mode

Ao

LA

Ao

Depth

12

LA

Distance

Time
Figure 1–11  3D, 2D, M-mode, and A-mode recordings of aortic valve motion. This illustration shows the following: the relationship between the 3D and
2D long-axis image of the aortic valve (left), which shows distance in both the vertical and horizontal direction; M-mode recording of aortic root (Ao), left atrium
(LA), and aortic valve motion, which shows depth-versus-time (middle); and A-mode recording (right), which shows depth only (with motion seen on the video
screen). Spatial relationships are best shown with 3D or 2D imaging, but temporal resolution is higher with M-mode and A-mode imaging.

from pulse to pulse, including the length of the ultrasound signal plus the time interval between signals.
Because a finite time is needed for each scan line
of data (depending on the depth of interest), the time
needed to acquire all the data for one image frame is
directly related to the number of scan lines and the
imaging depth. Thus, PRF is lower at greater imaging depths and higher at shallow depths. In addition,
there is a tradeoff between scan line density and
image frame rate (the number of images per second). For cardiac applications, a high frame rate (≥30
frames per second) is desirable for accurate display
of cardiac motion. This frame rate allows 33 ms per
frame or 128 scan lines per 2D image at a displayed
depth of 20 cm.
The reflected ultrasound signals for each scan line
are received by the piezoelectric crystal and a small
electric signal generated with:
  

n
n

 mplitude proportional to incident angle and
A
acoustic impedance
Timing proportional to distance from the
transducer

time-gain compensation (TGC), filtering (to reduce
noise), compression, and rectification. Envelope
detection generates a bright spot for each signal
along the scan line, which then undergoes analogto-digital scan conversion, since the original polar

coordinate data must be fit to a rectangular matrix
with appropriate interpolation for missing matrix
elements. This image is subject to further “postprocessing” to enhance the visual appreciation of tomographic anatomy and is displayed in “real time”
(nearly simultaneous with data acquisition) on the
monitor screen.
Although standard ultrasound imaging is based
on reflection of the fundamental transmitted frequency from tissue interfaces, tissue harmonic imaging
(THI) instead is based on the harmonic frequency
energy generated as the ultrasound signal propagates
through the tissues. These harmonic frequencies
result from the nonlinear effects of the interaction of
ultrasound with tissue and with the key properties:
  

n

  

n

This signal undergoes complex manipulation
to form the final image displayed on the monitor.
Typical processing includes signal amplification,

n
  

 armonic signal strength increases with depth
H
of propagation.

Harmonic frequencies are maximal at typical
cardiac imaging depths.
Stronger fundamental frequencies produce
stronger harmonics.


Principles of Echocardiographic Image Acquisition and Doppler Analysis  |  Chapter 1

Signal strength

Near the skin, very
few harmonics are
produced

n
Fundamental
Harmonics
At usual imaging
distances, harmonics
are much stronger

n

0

4

8

12

16

20

Distance [cm]
Figure 1–12  Relation between imaging distance and strength of fundamental and harmonic frequencies. As ultrasound pulse propagates, strength
of fundamental frequency declines, while strength of harmonic frequency increases. At usual imaging distances for cardiac structures, strength of harmonic frequency is maximized. In this schematic, harmonic frequency strength
is exaggerated; harmonic frequency signal strength is much lower than fundamental frequency signal strength.  (From Thomas JD, et al: Tissue harmonic
imaging: why does it work? J Am Soc Echocardiog 11:803-808, 1998.)

Thus, harmonic imaging reduces near-field and
side-lobe artifacts and improves endocardial definition,
particularly in patients with poor fundamental frequency images (Fig. 1-12). THI improves visualization
of the left ventricular (LV) endocardium, which allows
border tracing for calculation of ejection fraction,
reduces measurement variability, and results in visualization of more myocardial segments during stress
echocardiography. However, although THI improves
lateral resolution by 20-50%, it reduces axial resolution by 40 to 100%. Thus, valves and other planar
objects may appear thicker with harmonic, compared
to fundamental, frequency imaging, so that caution is
needed when diagnosing valve abnormalities or making measurements of chamber or vessel size.
Instrument Settings
Many of the elements in the process of image formation are features of a particular transducer and
instrument that cannot be modified by the operator.
However, for each patient and echocardiographic
view, optimal image quality depends on transducer
selection and instrument settings. Standard imaging
controls available in most ultrasound systems include:
  

n

n

 ower output: This control adjusts the total ultraP
sound energy delivered by the transducer in the
transmitted bursts; higher power outputs result
in higher-amplitude reflected signals (see Bioeffects and Safety, p. 27.).
Gain: Adjusts the displayed amplitude of the
received signals, similar to the volume control in
an audio system.

n

 GC: Allows differential adjustment of gain
T
along the length of the ultrasound beam to
compensate for the effects of attenuation. Nearfield gain can be set lower (because reflected
signals are stronger) with a gradually increased
gain over the midfield (“ramp” or “slope”) and
a higher gain in the far field (because reflected
signals are weaker). On some instruments, nearfield and far-field gain beyond the range of the
TGC are adjusted separately.
Depth: Displayed depth affects the PRF and
frame rate of the image and also allows maximal display of the area of interest on the screen.
Standard depth settings show the entire plane
(from the transducer down), while “resolution,”
“zoom,” or “magnification” modes focus on a
specific depth range of interest.

Dynamic range/compression: The amplitude range
(in dB) of the reflected signal is greater than the
display capacity of ultrasound systems so the
signal is compressed into a range of values from
white to black, or gray scale. The number of levels of gray in the image, or dynamic range, can be
adjusted to provide an image with marked contrast between light and dark areas or a gradation
of gray levels between the lightest and darkest
areas. A variation of standard gray scale is to use
color intensity for each amplitude value.

  

Other typical instrument controls include preprocessing and postprocessing settings that change the
appearance of the displayed image. Image quality and
resolution also depend on scan-line density and other
factors (see Table 1-4). Scan-line density (or frame rate
or both) can be increased by using a lower depth setting or by narrowing the sector to less than the standard 60° wide image.
Imaging Artifacts
Imaging artifacts include (1) extraneous ultrasound
signals that result in the appearance of “structures”
that are not actually present (at least at that location),
(2) failure to visualize structures that are present, and
(3) an image of a structure that differs in size or shape
or both from its actual appearance. Obviously, recognition of image artifacts is important for both the
individual performing the study and the individual
interpreting the echocardiographic data (Table 1-5).
The most common image “artifact” is suboptimal
image quality resulting from poor ultrasound tissue penetration related to the patient’s body habitus with interposition of high attenuation tissues (e.g., lung or bone)
or an increased distance (e.g., adipose tissue) between
the transducer and cardiac structures. While, strictly

speaking, poor image quality is not an “artifact,” a low
signal-to-noise ratio makes accurate diagnosis difficult
and precludes quantitative measurements. In many

13


14

Chapter 1  |  Principles of Echocardiographic Image Acquisition and Doppler Analysis

TABLE 1-5  Ultrasound Imaging Artifacts
Artifact

Mechanism

Example(s)

Suboptimal image
quality

Poor ultrasound tissue penetration

Body habitus (obesity, lung disease)
Postcardiac surgery

Acoustic shadowing

Reflection of entire ultrasound signal by a strong
specular reflector

Prosthetic valve
Calcification

Reverberations

Reverberation between two strong parallel reflectors

Prosthetic valve

Beam width

Superimposition of structures within the beam
profile (including side lobes) into a single
tomographic image

Aortic valve “in” LA
Atheroma “in” aortic lumen

Lateral resolution

Displayed width of a point target varies with depth

Excessive width of calcified mass or
prosthetic valve

Refraction

Deviation of ultrasound signal from a straight path
along the scan line

Double aortic valve or LV image in
short-axis view

Range ambiguity

Echo from previous pulse reaches transducer on
next cycle

Second, deeper heart image

Electronic
processing

Instrument specific

Variable

patients with suboptimal ultrasound penetration,
image quality is improved by use of tissue harmonic
imaging. In some cases, TEE imaging may be needed
to make an accurate diagnosis.
Acoustic shadowing (Fig. 1-13) occurs when a structure
with a marked difference in acoustic impedance (e.g.,
prosthetic valve, calcium) blocks transmission of the
ultrasound wave beyond that point. The image appears
devoid of reflected signals distal to this structure, since
no signal penetrates beyond the shadowing structure.
The shape of the shadow (like a light shadow) follows
the ultrasound path, so a small structure near the transducer casts a large shadow. When shadowing occurs,

an alternate acoustic window is needed for evaluation
of the area of interest. In some cases, a different transthoracic view will suffice. In other cases (e.g., prosthetic
mitral valve), TEE imaging may be necessary.
Reverberations (Fig. 1-14) are multiple linear highamplitude echo signals originating from two strong
specular reflectors resulting in back-and-forth reflection of the ultrasound signal before it returns to the
transducer. On the image, reverberations appear as
relatively parallel, irregular, dense lines extending
from the structure into the far field. Like acoustic
shadowing, prominent reverberations limit evaluation
of structures in the far field. In less dramatic cases,
reverberations may appear to represent abnormal
structures. For example, in the parasternal long-axis
view, a linear echo in the aortic root may originate as a
reverberation from anterior structures (e.g., ribs) rather
than representing a dissection flap.
The term beam width artifact is applied to two separate sources of image artifacts. First, remember that all

MVR
LA

S

R

S

Figure 1–13  Example of acoustic shadowing and reverberations.
TEE view in a patient with a valve replacement (MVR) shows shadowing (S)
by the sewing ring with reverberations (R) from the valve occluders further
obscuring the ventricle.

the structures within the 3D volume of the ultrasound
beam are displayed in a single tomographic plane. In
the focal zone of the beam, the 3D volume is quite
small and the tomographic “slice” is narrow. In the far
zone, however, strong reflectors at the edge of a larger
beam will be superimposed on structures in the central
zone of the beam even though signal intensity falls off
at the edges of the beam. In addition, strong reflectors in side lobes of the beam will be displayed in the


Principles of Echocardiographic Image Acquisition and Doppler Analysis  |  Chapter 1

LA

Transducer
Ao

RVOT

A B

Parallel
strong reflectors

Reverberations

A
B

Figure 1–15  Example of beam width artifact. Apparent “mass” attached to the aortic valve on this off-axis TEE view is the noncoronary cusp
of the aortic valve seen “en face.” Imaging in other planes demonstrated a
normal trileaflet aortic valve.

Ultrasound
artifacts

Figure 1–14  Reverberation artifacts result from the interaction of ultrasound with two parallel strong reflectors. The transmitted ultrasound
beam (red with down arrow) is reflected from the first reflector and returns
to the transducer (red with up arrow) resulting in an ultrasound signal that
corresponds to the correct depth of the reflector. However, ultrasound
signals also reflect back and forth between the two strong reflectors, with
some signals returning to the transducer after two (A), three (B), or more
reverberation cycles. The longer time from transmission to reception of
these late-returning signals results in their display on the ultrasound image at points distal to the actual reflector. In clinical imaging, reverberation
artifacts can either appear as a single linear signal distal to the actual object
or as a band of signals obscuring distal structures (see Fig. 1-13) because
of multiple parallel reflectors.

tomographic section corresponding to the main beam
(Fig. 1-15).
The second type of beam width artifact is a consequence of varying lateral resolution at different
imaging depths. A point target appears as a line whose
length depends on the beam characteristics at that
depth and the amplitude of the reflected signal. For
example, the struts on a prosthetic valve can appear
much longer than their actual dimension because of
poor lateral resolution. Sometimes beam width artifacts can be mistaken for abnormal structures such as a
valvular vegetation, an intracardiac mass, or an aortic
dissection flap.

The appearance of a side-by-side double image
results from ultrasound refraction as it passes through
a tissue proximal to the structure of interest. This
artifact often is seen in parasternal short-axis views

of the aortic valve or LV, where a second valve or LV
is “seen” medial to and partly overlapping the actual
valve or LV. The explanation for this appearance
is that the transmitted ultrasound beam is deviated
from a straight path (the scan line) by refraction as it
passes through a tissue near the transducer. When this
refracted beam is reflected back to the transducer by a
tissue interface, the reflected signal is assumed to have
originated from the scan line of the transmitted pulse
(Fig. 1-16) and thus is displayed on the image in the
wrong location.
Range ambiguity occurs when echo signals from an
earlier pulse cycle reach the transducer on the next
“listen cycle” for that scan line, resulting in deep structures appearing closer to the transducer than their
actual location. The appearance of an anatomically
unexpected echo within a cardiac chamber often is
due to range ambiguity, as can be demonstrated by the
disappearance or a change in position of this artifact
when the depth setting (and PRF) is changed. Another
type of range ambiguity is the appearance of an
apparent second heart, deeper than the actual heart—
a double image on the vertical axis. This type of range
ambiguity results from echoes being re-reflected by a
structure close to the transducer (such as a rib), being
re-reflected by the cardiac structures and thus received

at the transducer at a time twice normal. This artifact
can be eliminated (or obscured) by decreasing the
depth setting or adjusting the transducer position to a
better acoustic window.

15