Feigenbaum's Echocardiography, 6th Edition PDF

FEIGENBAUM’S ECHOCARDIOGRAPHY th 6 EDITION THIS BOOK BELONGS TO JUSTIN JOSE TABLE OF CONTENTS Front Cover 3 Editors...

FEIGENBAUM’S ECHOCARDIOGRAPHY th 6 EDITION THIS BOOK BELONGS TO JUSTIN JOSE TABLE OF CONTENTS Front Cover 3 Editors 4 DEDICATION 5 PREFACE 6 ACKNOWLEDGMENTS 8 1 - History of Echocardiography 9 2 - Physics and Instrumentation 33 3 - Specialized Echocardiographic Techniques and Methods 110 4 - Contrast Echocardiography 174 5 - The Echocardiographic Examination 248 6 - Evaluation of Systolic and Diastolic Function of the Left Ventricle 309 7 - Left Atrium, Right Atrium, and Right Ventricle 403 8 - Hemodynamics 467 9 - Pericardial Diseases 534 10 - Aortic Valve Disease 596 11 - Mitral Valve Disease 649 12 - Tricuspid and Pulmonary Valves 758 13 - Infective Endocarditis 817 14 - Prosthetic Valves 867 15 - Coronary Artery Disease a - CLINICAL OVERVIEW 937 b - DETECTION AND QUANTITATION OF WALL MOTION ABNORMALITIES 951 c - ECHOCARDIOGRAPHIC EVALUATION OF CLINICAL SYNDROMES 972 d - COMPLICATIONS OF ACUTE MYOCARDIAL INFARCTION 994 e - CHRONIC CORONARY ARTERY DISEASE 1016 f - DIRECT CORONARY VISUALIZATION 1039 g - SUGGESTED READINGS 1054 16 - Stress Echocardiography 1062 17 - Cardiomyopathies 1134 18 - Congenital Heart Diseases a - INTRODUCTION 1228 b - THE ECHOCARDIOGRAPHIC EXAMINATION - A SEGMENTAL APPROACH TO ANATOMY 1230 c - ABNORMALITIES OF RIGHT VENTRICULAR INFLOW 1239 d - ABNORMALITIES OF LEFT VENTRICULAR INFLOW 1244 e - ABNORMALITIES OF RIGHT VENTRICULAR OUTFLOW 1255 f - ABNORMALITIES OF LEFT VENTRICULAR OUTFLOW 1266 g - COARCTATION OF THE AORTA 1277 h - ABNORMALITIES OF CARDIAC SEPTATION 1282 i - ABNORMAL VASCULAR CONNECTIONS AND STRUCTURES 1328 j - CONOTRUNCAL ABNORMALITIES 1343 k - ABNORMALITIES OF VENTRICULAR DEVELOPMENT 1365 l - ECHOCARDIOGRAPHIC EVALUATION DURING AND AFTER SURGERY 1373  m - FETAL ECHOCARDIOGRAPHY 1381 n - SUGGESTED READINGS 1387 19 - ICU and Operative, Perioperative Applications 1395 20 - Diseases of the Aorta 1481 21 - Masses, Tumors, and Source of Embolus 1553 22 - Echocardiography in Systemic Disease and Clinical Problem Solving 1626 Editors: Feigenbaum, Harvey; Armstrong, William F.; Ryan, Thomas Title: Feigenbaum's Echocardiography, 6th Edition Copyright ©2005 Lippincott Williams & Wilkins > Front of Book > Editors Editors Harvey Feigenbaum MD Distinguished Professor of Medicine Indiana University School of Medicine, Krannert Institute of Cardiology, Indianapolis, Indiana William F. Armstrong MD Professor of Medicine; Director Echocardiography Laboratory, University of Michigan, Ann Arbor, Michigan Thomas Ryan MD Professor of Medicine; Director Duke Heart Center, Duke University Health System, Durham, North Carolina Secondary Editors Ruth Weinberg Acquisitions Editor Joyce Murphy Developmental Editor Alicia Jackson Production Editor Benjamin Rivera Manufacturing Manager Kathy Neely Marketing Manager Larry Didona Cover Designer Compositor: Graphic World, Inc. Printer: Quebecor World Taunton Editors: Feigenbaum, Harvey; Armstrong, William F.; Ryan, Thomas Title: Feigenbaum's Echocardiography, 6th Edition Copyright ©2005 Lippincott Williams & Wilkins > Front of Book > DEDICATION DEDICATION To our children and grandchildren, who make it all worthwhile - Steve, Tom, Lyle, Andrew, Russell, Megan, Katie, Patrick, Tess, Olivia, Jake, and Lucy Editors: Feigenbaum, Harvey; Armstrong, William F.; Ryan, Thomas Title: Feigenbaum's Echocardiography, 6th Edition Copyright ©2005 Lippincott Williams & Wilkins > Front of Book > PREFACE PREFACE This latest edition of Echocardiography represents a major transition not only in the field of echocardiography but in the history of this publication. As with previous editions there have been numerous important advances which warrant an updated review of the field. A more fundamental change is that echocardiography has now become the backbone of cardiac imaging. It is an integral part of clinical cardiology. With the introduction of digital acquisition and display of echocardiograms, the images and reports have become ubiquitous, with clinicians having access to the recordings throughout the clinical environment. It is now possible to access reports and images of echocardiograms from almost anywhere including the emergency room, office, outpatient facility and even from home or a hotel. There are very few situations in clinical cardiology where an echocardiogram is not of significant value. Therefore virtually all physicians who care for cardiology patients have to be familiar with echocardiography. Furthermore, the description and understanding of cardiologic problems almost always involves the echocardiographic findings. As a result of this fundamental change in the role of echocardiography, the sixth edition is much more clinical in nature. A brief review of the clinical problem in which echocardiography is useful is provided in every chapter. How echocardiography integrates with clinical practice is emphasized far more than it has been in previous editions. In past editions a large percentage of the illustrations were borrowed from the literature. This edition represents a major departure from this practice. First of all it is far more difficult to reproduce illustrations from prior publications. But more importantly, virtually all of the illustrations in this edition were acquired digitally. This practice enhances the quality of the illustrations. More significantly, it provides a real-time moving display of the two-dimensional and color Doppler images. Thus, a DVD-ROM with extensive moving images is provided with this text. Since echocardiography is an imaging technique, an exhaustive number of illustrations is essential for a proper understanding of this diagnostic modality. Probably the most important deviation from prior editions is that the sixth edition is no longer principally a single-author textbook. The first three editions of the book were entirely written by me. However, the fourth edition had a chapter written by William Armstrong and the fifth edition had a chapter by Tom Ryan. This sixth edition is being written by all three of us. The bulk of the writing, with rare exception, has been by Bill and Tom. This change in authorship was done for various reasons. First of all it is almost impossible now for a single author to cover the entire field of echocardiography satisfactorily. The developments in the field have been too numerous, and the role that echocardiography plays in clinical cardiology requires almost a mini cardiology textbook. Furthermore this text represents an example of a “changing of the guard.†In my judgment Bill and Tom represent the next generation of senior authorities in this field. Despite the fact that this book technically is a multi-author textbook, we have made a great effort for it to read as if it were written by a single person. This feature has always been important to me because I feel that it is a much better educational experience for the reader if all of the chapters are integrated and written in a similar style. Since both Bill and Tom have contributed to previous editions, it is natural for them to be the principal authors of this newest edition. All three of us have read every chapter and made editorial comments. We have made every effort for the chapters to be similar so that one cannot tell who actually wrote the chapter. Bill and Tom began their careers in echocardiography at Indiana University, so our understanding of the technique has the same basic foundation. The other major necessity for this text to come from multiple institutions is the fact that since almost all of the illustrations are original, it would be very difficult for any one institution to have appropriate illustrations for everything in the field of echocardiography. Thus the illustrations come from all three institutions and cover the field extremely well. In past editions I made an effort for the references to be exhaustive so that the textbook could be used as a reference library. In this day of the Internet and ready access to multiple references, this need is no longer present. As a result the references are not nearly as extensive, but we attempt to give the reader some guidance as to where more references can be found in the literature. Harvey Feigenbaum MD Editors: Feigenbaum, Harvey; Armstrong, William F.; Ryan, Thomas Title: Feigenbaum's Echocardiography, 6th Edition Copyright ©2005 Lippincott Williams & Wilkins > Front of Book > ACKNOWLEDGMENTS ACKNOWLEDGMENTS A project such as this would be impossible without the assistance and support of many people. Although it is impractical to thank them all individually, the authors wish to express their gratitude to the sonographers, fellows, and colleagues who contributed ideas, suggestions, and illustrations. We especially want to acknowledge Allison Thodoroff and David Adams for their superb help with the text and illustrations. Among other things, Allison and David created most of the illustrations and digital loops for the text. Their contribution to the final product is immeasurable. Several colleagues read portions of the text and made substantive suggestions, particularly Sidney Edelman, PhD, who provided extremely helpful suggestions in the chapter on physics and instrumentation. Finally, the secretarial support of Cheryl Childress, Karen Spirl, Joyce Price, Hope Odzak, and Pam King is warmly recognized. Editors: Feigenbaum, Harvey; Armstrong, William F.; Ryan, Thomas Title: Feigenbaum's Echocardiography, 6th Edition Copyright ©2005 Lippincott Williams & Wilkins > Table of Contents > Chapter 1 - History of Echocardiography Chapter 1 History of Echocardiography Harvey Feigenbaum M.D. Many histories of diagnostic ultrasound, and cardiac ultrasound in particular, have been written.1 , 2 , 3 , 4 , 5 , 6 They all seem to address this field from a different perspective. One can begin the history in the twentieth century, Roman times, or any of the centuries in between. It is stated that a Roman architect, Vitruvius, first coined the word echo.7 A Franciscan friar, Marin Mersenne (1588–1648), is frequently called the “father of acoustics†because he first measured the velocity of sound.7 Another early physicist, Robert Boyle (1627–1691), recognized that a medium was necessary for the propagation of sound.7 Abbe Lazzaro Spallanzani (1727–1799) is frequently referred to as the “father of ultrasound.†8 H e demonstrated that bats were blind and in fact navigated by means of echo reflection using inaudible sound. In 1842, Christian Johann Doppler (1803–1853) noted that the pitch of a sound wave varied if the source of the sound was moving.9 He worked out the mathematical relationship between the pitch and the relative motion of the source and the observer. The ability to create ultrasonic waves came in 1880 with the discovery of piezoelectricity by Curie and Curie. 1 0 , 1 1 They noted that if certain crystalline materials are compressed, an electric charge is produced between the opposite surfaces. They then noted that the reverse was also true. If an electrical potential is applied to a crystal, it is compressed and decompressed depending on the polarity of the electric charge, and thus very high frequency sound can be produced. In 1912, a British engineer, L. F. Richardson, suggested that an echo technique could be used to de- tect underwater objects. Later during World War I, Paul Langevin was given the duty of detecting enemy submarines using sound, which culminated in the development of sonar.3 Sokolov1 2 described a method for using reflected sound to detect metal flaws in 1929. In 1942, Floyd Firestone,1 3 an American engineer, began to apply this technique and received a patent. It is this flaw detection technique that ultimately was used in medicine. An Austrian, Karl Dussik,1 4 was probably the first to apply ultrasound for medical diagnosis in 1941. He initially attempted to outline the ventricles of the brain. His approach used transmission ultrasound rather than reflected ultrasound. After World War II, many of the technologies developed during that war, including sonar, were applied for peaceful and medical uses. In 1950, W. D. Keidel,1 5 a German investigator, used ultrasound to examine the heart. His technique was to transmit ultrasonic waves through the heart and record the effect of ultrasound on the other side of the chest. The purpose of his work was to try to determine cardiac volumes. The first effort to use pulse- reflected ultrasound, as described by Firestone, to examine the heart was initiated by Dr. Helmut Hertz of Sweden. He was familiar with Firestone's observations and in 1953 obtained a commercial ultrasonoscope, which was being used for nondestructive testing. He then collaborated with Dr. Inge Edler who was a practicing cardiologist in Lund, Sweden. The two of them began to use this commercial ultrasonoscope to examine the heart. This collaboration is commonly accepted as the beginning of clinical echocardiography as we know it today.1 6 The original instrument (Fig. 1.1 ) was quite insensitive. The only cardiac structures that they could record initially were from the back wall of the heart. In retrospect, these echoes probably came from the posterior left ventricular wall. With some modification of their instrument, they were able to record an echo from the anterior leaflet of the mitral valve. However, they did not recognize the source of this echo for several years and originally attributed the signal to the anterior left atrial wall. Only after some autopsy investigations did they recognize the echo's true origin. Edler1 7 went on to perform a number of ultrasonic studies of the heart. Many of the cardiac echoes currently used were first described by him. However, the principal clinical application of echocardiography developed by Edler was the detection of mitral stenosis.1 8 He noted that there was a difference between the pattern of motion of the anterior mitral leaflet in patients who did or did not have mitral stenosis. Thus, the early P.2 studies published in the mid-1950s and early 1960s primarily dealt with the detection of this disorder.  FIGURE 1.1. Ultrasonoscope initially used by Edler and Hertz for recording their early echocardiograms. (From Edler I, Ultrasoundcardiography. Acta Med Scand Suppl 370 1961;170:39, with permission.) The work being done in Sweden was duplicated by a group in Germany headed by Dr. Sven Effert.1 9 , 2 0 Their publications began to appear in the late 1950s and were primarily duplications of Edler's work describing mitral stenosis. One notable observation made by Effert and his group2 0 was the detection of left atrial masses. Schmitt and Braun2 1 in Germany also began working with ultrasound cardiography and published their work in 1958, again repeating what Edler and Effert had been doing. Edler and his co-workers2 2 developed a scientific film that was shown at the Third European Congress of Cardiology in Rome in 1960 2 2. Edler et al. 2 3 also wrote a large review of cardiac ultrasound as a supplement to Acta Medica Scandinavica , which was published in 1961, and remained the most comprehensive review of this field for more than 10 years. In the movie and the review, Edler and his co-investigators described the ultrasonic techniques for the detection of mitral stenosis, left atrial tumors, aortic stenosis, and anterior pericardial effusion. Despite their initial efforts at using ultrasound to examine the heart, neither Edler nor Hertz really anticipated that this technique would flourish. Helmut Hertz was primarily interested in being able to record the ultrasonic signals. In the process, he developed ink jet technology and only spent a few years in the field of cardiac ultrasound. He devoted most of the rest of his career to ink jet technology, for which he held many important patents. He also advised Siemens Corporation, who provided their first ultrasonic instrument, that they should not enter the field of cardiac ultrasound because he personally did not feel that there was a great future in this area (Effert, personal communication, 1996). Edler too did not develop any further techniques in cardiac ultrasound. He retired in 1976 and until then was primarily concerned with the application of echocardiography for mitral stenosis and, to a lesser extent, mitral regurgitation. He never became involved with any of the newer techniques for pericardial effusion or ventricular function. China was another country where cardiac ultrasound was used in the early years. In the early 1960s, investigators both in Shanghai and Wuhan were using ultrasonic devices to examine the heart. They began initially with an A-mode ultrasound device and then developed an M-mode recorder. 2 4 , 2 5 The investigators duplicated the findings of Edler and Effert with regard to mitral stenosis.2 6 Unique contributions of the Chinese investigators included fetal echocardiography2 7 and contrast echocardiography using hydrogen peroxide and then carbon dioxide.2 8 In the United States, echocardiography was introduced by John J. Wild, HD Crawford, and John Reid2 9 who examined the excised heart. They were able to identify a myocardial infarction and published their findings in 1957 in the American Heart Journal. Neither Wild nor Reid was a physician. Reid was an engineer who subsequently went to the University of Pennsylvania for his doctorate degree. While there, he wanted to continue his interest in examining the heart ultrasonically. He joined forces with Claude Joyner, who was a practicing cardiologist in Philadelphia. Reid proceeded to build an ultrasonoscope, and Joyner and he began duplicating the work on mitral stenosis that was described by Edler and Effert. This work was published in Circulation in 1963 and represents the first American clinical effort using pulsed reflected ultrasound to examine the heart.3 0 I became interested in echocardiography in the latter part of 1963. While operating a hemodynamic laboratory and becoming frustrated with the limitations of cardiac catheterization and angiography, I saw an ad from a now defunct company that was claiming that it had an instrument that could measure cardiac volumes with ultrasound. This claim ultimately proved to have no basis. However, when I first saw the ultrasound instrument displayed at the American Heart Association meeting in Los Angeles in 1963, I placed the transducer on my chest and saw a moving echo, which had to be coming from the posterior wall of my heart. This signal undoubtedly was the same echo that Hertz and Edler had noted approximately 10 years earlier. I had the people from the company explain the principles by which such a signal might be generated. I asked them whether fluid in back of the heart would give a different type of a signal, and they said that fluid would be echo free. When I returned to Indiana, I found that the neurologists had an ultrasonoscope that they used for detecting the midline of the brain. Fortunately for me, the instrument was rarely being used and I was able to borrow it. I proceeded to examine more individuals, and again I was able to record an echo from the back wall of the left ventricle. I looked for a patient with P.3 pericardial effusion. As predicted, there were now two echoes separated by an echo-free space. The more posterior echo no longer moved, whereas the more anterior echo moved with cardiac motion. We went to the animal laboratory to confirm these findings and thus began my personal career in cardiac ultrasound. This initial paper on pericardial effusion was published in the Journal of the American Medical Association in 1965.3 1 Although this phase of the history of echocardiography is commonly considered the origins of the early practice of echocardiography, it should be mentioned that Japanese investigators were working simultaneously using ultrasound to examine the heart. In the mid- 1950s, several Japanese investigators such as Satomura, Yoshida, and Nimura at Osaka University were using Doppler technology to examine the heart. They began publishing their work in the mid-1950s.3 2 , 3 3 These efforts laid the basis for much of what we do today with Doppler ultrasound. The field of cardiac ultrasound has evolved with the efforts of numerous individuals over the past 50 years. This development is an outstanding example of collaboration between physicists, engineers, and clinicians. Each of the cardiac ultrasonic techniques has its own individual history. Even the name echocardiography has a story of its own. Edler and Hertz first called this technique ultrasound cardiography with the abbreviation being UCG. Ultrasound cardiography was a somewhat cumbersome name. The most common use of medical diagnostic ultrasound in the late 1950s and early 1960s was an A-mode technique to detect the midline of the brain. This midline echo would shift if there were an intracranial mass. The technique was known as echoencephalography, and the instrument was an echoencephalograph. It was such an instrument that I borrowed from the neurologists. If the ultrasonic examination of the brain is echoencephalography, then an examination of the heart should be echocardiography. Unfortunately, the abbreviation for an echocardiogram would be ECG, which was already preempted by electrocardiography. We could not use the abbreviation “echo†because it did not differentiate from an echoencephalogram. The reason the term echocardiography finally caught on was because echoencephalography disappeared. No other diagnostic ultrasound technique used the term echo except for the examination of the heart. So the abbreviation echo†now stands only for echocardiography and is not confused with any other ultrasonic examination. DEVELOPMENT OF VARIOUS ECHOCARDIOGRAPHIC TECHNOLOGIES The story of echocardiography involves the evolution and development of its many modalities such as A-mode, M-mode, contrast, two- dimensional, Doppler, transesophageal, and intravascular applications. The Doppler story is truly lengthy and international. The Japanese began working with Doppler ultrasound in the mid-1950s.3 2 , 3 3 American workers, such as Robert Rushmer in Seattle, were early investigators using Doppler techniques.3 4 Dr. Rushmer was a recognized expert in cardiac physiology. John Reid later moved to Seattle and joined Rushmer and his group in developing Doppler technology. One of the engineers, Donald Baker, was in that group and developed one of the first pulsed Doppler instruments. 3 5 Eugene Strandness was a vascular surgeon in Seattle using Doppler for peripheral arterial disease.3 6 European investigators were also very active in using Doppler technology. Several early French workers, namely Peronneau3 7 and later Kalmanson, 3 8 wrote extensively on the use of Doppler ultrasound to examine the cardiovascular system. A major development in Doppler ultrasound came when Holen3 9 and then Hatle4 0 demonstrated that one could derive hemodynamic information from Doppler ultrasound. They noted that one could use a modified version of the Bernoulli equation to detect gradients across stenotic valves. The report that the pressure gradient of aortic stenosis could be determined with Doppler ultrasound was probably the development that established Doppler echocardiography as a clinically important technique. The field of contrast echocardiography began with an unexpected observation by Gramiak et al.4 1 at the University of Rochester. They apparently were doing an ultrasonic examination on a patient undergoing an indicator dilution test using indocyanine green dye. Much to their surprise, they noticed a cloud of echoes introduced into the cardiovascular system with the injection of dye. Apparently, Joyner had noticed a similar observation with the injection of saline but did not report the finding. I heard Gramiak present his group's work at a meeting and promptly used that technique to help establish the echocardiographic identity of the left ventricular cavity. 4 2 Workers at the Mayo Clinic headed by Jamil Tajik and Jim Seward went on to use this contrast technique in a very eloquent way to identify right-to-left shunts.4 3 Contrast agents have evolved to the current commercial products, which are manufactured. The tiny echo-producing bubbles are small enough to pass through capillaries so that a peripheral injection can be seen on the left side of the heart. 4 4 Two-dimensional echocardiography has a lengthy and fascinating history. As with almost every aspect of cardiac ultrasound, there is an international flavor to this story. Two-dimensional ultrasonic scanning dates back to early workers such as Douglass Howry when he began using compound scanning for various parts of the body. One of his early compound scanners used a transducer that was mounted on a ring from a B29 gun turret.4 5 The Japanese introduced a variety of ultrasonic devices to create two-dimensional recordings of the heart.4 6 They used elaborate water baths and scanning techniques (Fig. 1.2 ). Gramiak and co-workers4 7 at the University of Rochester used reconstructive two-dimensional M-mode techniques P.4 to create ultrasonic “cinematography†(Fig. 1.3 ). Donald King4 9 in New York developed a stop-action type of technique for creating a reconstructed two-dimensional image of the heart (Fig. 1.4 ). FIGURE 1.2. Relatively early system using a mechanical sector scanner and a water bath to obtain cross-sectional echograms of the heart. (From Ebina T, Oka S, Tanaka N, et al. The ultrasono- tomography of the heart and the great vessels in living human subjects by means of the ultrasonic reflection technique. Jpn Heart J 1967;8:331, with permission.) A major breakthrough occurred when an engineer, Nicholas Bom, in Rotterdam, developed a linear scanner (Fig. 1.5 ).4 9 By using multiple crystals, he could create a rectangular image of the heart in real time. Although this technique ultimately never proved to be useful in examining the heart, partially because of the rib shadows, this technique did show the virtue of real-time imaging. It ultimately proved to be a leading form of two-dimensional imaging in other parts of the body but not the heart. Real-time two-dimensional echocardiography became practical by using a sector scan rather than a linear scan. Initially, the scan devices were mechanical. Griffith and Henry5 0 at the National Institutes of Health developed a mechanical device that rocked the transducer back and forth. The device was handheld; however, the ability to manipulate the transducer was very limited. Reggie Eggleton, who originally worked at the University of Illinois with Robert, Frank, and Elizabeth Frye, moved to Indiana and developed a mechanical two-dimensional scanner (Fig. 1.6 ). Interestingly enough, his first prototype was actually a modified Sunbeam electric toothbrush. This early mechanical scanner was the first commercially successful real-time two-dimensional device.5 1 Eventually, mechanical sector scanners were replaced by phased-array technology, which was initially developed by Fritz Thurstone and Olaf vonRamm at Duke University.5 2 FIGURE 1.3. Frames from a movie film using a spatially oriented reconstruction of the M-mode echogram to produce a pseudo real- time, cross-sectional examination of the mitral valve motion. The two enlarged frames show the position of the mitral valve (arrow) in systole and diastole. (From Gramiak R, Waag R, Simon W. Cine ultrasound cardiography. Radiology 1973;107:175, with permission.) FIGURE 1.4. Compound, electrocardiograph-gated cross-sectional examination of the heart. RVO, right ventricular outflow tract; AW, anterior wall of the aorta; AV, aortic valve; PW, posterior wall of the aorta; LA, left atrium; VS, interventricular septum; AMV, anterior mitral valve leaflet; LV, left ventricle; CW, chest wall. (From King DL, Steeg CN, Ellis K. Visualization of ventricular septal defect by cardiac ultrasonography. Circulation 1973;48:1215, with permission.)  FIGURE 1.5. Photograph of a multielement transducer that provides an electronic linear scan of the heart. This probe consists of 20 individual piezoelectric elements. (From Bom N, Lancee CT, Van Zwienten G, et al. Ultrascan echocardiography. I. Technical description. Circulation 1973;48:1066, with permission.) P.5 Color flow Doppler or two-dimensional Doppler ultrasound dates back to the late 1970s. A group headed by Brandestini working at the University of Washington in Seattle showed how one could use an M- mode recording of a multigated Doppler signal (Fig. 1.7 ). 5 3 They encoded the Doppler signal with color to indicate the direction of flow. This principle was later more fully developed by Japanese workers including Kasai et al.5 4 The key to the development of their two- dimensional color display was the autocorrelation detection of the Doppler velocities. They were now able to provide an excellent real- time two-dimensional display of color flow. Omoto, a Japanese cardiovascular surgeon, and co-workers5 5 helped to popularize the clinical value of two-dimensional color Doppler imaging. FIGURE 1.6. Photograph of a hand-held mechanical sector scanner. (From Eggleton RC, Feigenbaum H, Johnston KW, et al. Visualization of cardiac dynamics with real-time B-mode ultrasonic scanner. In: White D, ed. Ultrasound in Medicine. New York: Plenum Publishing, 1975:1385, with permission.) FIGURE 1.7. Combined M-mode and Doppler recording whereby the Doppler signal is superimposed on the M-mode tracing. The direction and velocity of the Doppler signal are displayed in varying colors. This particular recording shows the right ventricular outflow tract (RVOT) and aorta. (From Brandestini MA, Eyer MK, and Stevenson JG. M/Q: M/Q-mode echocardiography. The synthesis of conventional echo with digital multigate Doppler. In: Lancee CT,  ed. Echocardiography. The Hague, Netherlands: Martinus-Nijhoff, 1979, with permission.) The origin of transesophageal echocardiography also dates back to the 1970s. Lee Frazin, a cardiologist in Chicago, placed an M-mode transducer at the tip of a transesophageal probe and demonstrated how one could obtain an M-mode recording of the heart via the esophagus.5 6 This technique never became clinically popular. However, both Japanese and European investigators began working with this technology.5 7 , 5 8 They all attempted to obtain two-dimensional images with a transesophageal probe. Initially, the devices were mechanical and later became electronic. Hisanaga and co-workers5 7 were among the Japanese engineers, and Jacques Souquet was a European engineer who made a major contribution to transesophageal electronic probes in 1982.5 9 Most of the early clinicians who demonstrated the utility of transesophageal echocardiography were European. The versatility of ultrasound is exemplified by the fact that one can devise ultrasonic imaging techniques using very large or very small transducers. An exquisite ultrasonic imaging device used to examine the entire body was developed by an Australian engineer, George Kossoff. He developed an instrument called an Octoson. It P.6 consisted of eight very large transducers that rotated around the body. The instrument produced images that were of excellent resolution and clarity. The other extreme is the ability to put a tiny transducer on the tip of a catheter that can be inserted in the cardiovascular system. Reggie Eggleton devised a catheter-based imaging system in the 1960s as did Ciezynski in Europe and Omoto in Japan. In the early 1970s, Nicholas Bom and colleagues6 0 described a real-time intracardiac scanner using a circular array of 32 elements at the tip of a catheter. This technology developed further to the point that catheter-tipped transducers could be placed on an intracoronary device. Such instruments have been used clinically and for investigational purposes for many years now. Possibly the clinician who used intracoronary ultrasound to its greatest extent is Steven Nissen, who currently is at the Cleveland Clinic. He has used this technique to revolutionize our understanding of coronary atherosclerosis.6 1 There has been interest in three-dimensional echocardiography for many years. Numerous efforts at using compound two-dimensional scans to produce three-dimensional imaging have been demonstrated.6 2 , 6 3 Some of these compound three-dimensional devices have been used clinically. There has been active investigation using real-time three-dimensional ultrasound. Among the leaders in this effort is Olaf vonRamm and his group.6 4 Handheld echocardiographs date back to 1978.6 5 This early device did not have sufficient image quality to be useful. However, now several such instruments are available and increasing in popularity. RECORDING ECHOCARDIOGRAMS Along with developing instruments to create images and physiologic information of the heart, there has been a simultaneous history of developing techniques for recording this information. From the very beginning, Helmut Hertz was primarily interested in recording rather than creating ultrasonic images. In so doing, he developed ink jet technology, which proved to be extremely important. When I first began using ultrasound in the early 1960s, a Polaroid camera was the principal recording technique for A-mode and M-mode echocardiograms (Figs. 1.8 and 1.9 ). This approach was extremely limited and had many problems. Some investigators, such as Gramiak, used 35 mm film to record their M-mode echocardiograms. Much of my early efforts were to get commercial companies to provide strip chart recorders for our M-mode echocardiograms. The variety of strip chart recorders that became available has its own history. With the advent of two- dimensional echocardiography, we had to work out a scheme for recording these real-time two-dimensional images. At our own institution, we first used super 8 movie film as our recording medium. We would direct a movie camera at the oscilloscope and generate movies. The use of movie film was short-lived and we soon went to videotape. Initially, we used reel-to-reel tape recorders. Then a variety of recorders with cassettes became available. A popular tape recorder in the early years was produced by Sanyo. Unfortunately, analyzing a study frame by frame was very tedious. One had to turn a small button-like control and could not view images backward. Finally, Panasonic developed a tape recorder that permitted easy forward and backward viewing as well as frame-by-frame analysis.  FIGURE 1.8. Early M-mode echocardiograph using a Polaroid camera to record an echocardiogram. Because of the dominance of two-dimensional echocardiography in the clinical use of echocardiography, videotape became the standard means of recording echocardiogram for decades. Unfortunately, videotape also has major limitations. Looking at serial studies with videotape is problematic. The accessibility of videotape is inconvenient. One cannot make measurements from videotaped images. Copies of videotaped images are always degraded. Digital recording of echocardiograms began in the early 1980s. Interest in using digital techniques has been accelerating ever since. There are numerous advantages to using a digital recording. Side-by-side comparisons are facilitated. One can make measurements easily, and the images are more accessible. Initially, the digital images were generated by grabbing the video signal either from the instrument or by digitizing the videotape. In recent years, a direct digital output from ultrasonic instruments has become available. Digital recording standards using DICOM have facilitated the use of digital imaging and have become a major factor in the general utility of this approach.  FIGURE 1.9. Early Polaroid recordings of M-mode echocardiogram. A : Mitral stenosis, B : normal mitral valve, C : pericardial effusion, D : dilated nonmoving left ventricle. P.7 CARDIAC SONOGRAPHERS Early in my experience with cardiac ultrasound, it became apparent that the technique would become fairly popular. Performing the echocardiograms myself became a fairly time-consuming activity. Being a clinical cardiologist with responsibilities for patient care, including cardiac catheterization, I clearly felt that I could not continue to be the principal person to obtain echocardiograms. We also did not have sufficient physicians interested in the technique to provide a complement of physicians to do the echocardiograms throughout the day. As a result, I believed that it would be possible to train a nonphysician to do an echocardiogram. There was considerable P.8 skepticism among the few physicians active in the field of ultrasound at the time as to whether this approach was feasible. The first nonphysician hired to perform echocardiograms was Charles Haine. Our second cardiac sonographer was Sonia Chang. Her skills in obtaining an M-mode echocardiogram were so outstanding that with my encouragement she eventually published a book on the M-mode echocardiographic examination. It was a major publication from which many of the early users of M-mode echocardiography learned their technical skills. Most of the visitors who came to Indiana in the early days learned how to do echocardiograms from Sonia. Sonia left Indiana just after the introduction of two-dimensional echocardiography. She went to Emory University in Atlanta to work with Dr. Willis Hurst, who was the Chairman of Cardiology at the time. Virtually every echocardiographic laboratory in the United States has a sonographer who excels in the ability to obtain an echocardiogram. Cardiac sonographers have been a major factor in making echocardiography a cost-effective examination. Using a nonphysician to create echocardiograms is not a worldwide concept. In most countries, echocardiograms are still obtained by physicians. P.9 One exception is England, where there is a somewhat different situation. Their cardiac sonographers are probably more highly trained individuals than our sonographers. They come closer to being a physician's assistant and have a greater formal education in cardiac physiology and anatomy. They also perform interpretations with a higher degree of frequency than do sonographers in the United States.  FIGURE 1.10. The program for the first course devoted to diagnostic ultrasound and cardiovascular disease held in Indianapolis in January 1968. ECHOCARDIOGRAPHIC EDUCATION AND ORGANIZATIONS The first meeting dedicated solely to cardiac ultrasound was in Indianapolis in January 1968 (Fig. 1.10 ). Among the faculty were Drs. Edler, Joyner, Reid, and Strandness (Fig. 1.11 ). There were approximately 50 people who attended that course, one of whom was Raymond Gramiak. At that meeting, Dr. Edler showed the movie that he had created for the 1960 European Congress of Cardiology Meeting in Rome. Another member of the faculty was Richard Popp, who was a cardiology fellow at Indiana at the time. Bernard Ostrum, who was a radiologist at Albert Einstein Medical Center, presented data on abdominal aortas. Chuck Haine was an integral part of the program and demonstrated some of our ultrasonic techniques at Indiana. FIGURE 1.11. Photograph of Drs. Edler and Feigenbaum demonstrating an M-mode echocardiograph at the 1968 meeting of cardiac ultrasound in Indianapolis. The American Society of Echocardiography was also created in Indianapolis in 1975. The decision to create the society was made at a postgraduate meeting in Indianapolis. The Journal of the American Society of Echocardiography began in 1988 and the first annual American Society of Echocardiography scientific meeting was held in Washington, DC in 1990. There are now several worldwide echocardiography organizations, publications, and meetings. Echocardiography has come a long way since its beginnings in the mid- 1950s. Although there are many new, highly sophisticated imaging technologies being developed, there is every reason to believe that the clinical utility and popularity of echocardiography will continue to grow. This diagnostic tool is amazingly versatile. It is still very cost- effective compared with competing technologies and has many new possibilities as to how this examination can be improved and provide more and better information. Thus, the future of echocardiography should be as productive and exciting as have been the previous five decades. REFERENCES 1. Feigenbaum H. Echocardiography. 1st Ed. Philadelphia: Lea & Febiger, 1972. 2. Holmes JH. Diagnostic ultrasound during the early years of A.I.U.M. J Clin Ultrasound 1980;8:299–308. 3. Wild PW. Early history of echocardiography. J Cardiovasc Ultrasonogr 1996;5:2. 4. Goldberg P, Kimmelman BA. Medical Diagnostic Ultrasound: A Retrospective on its 40th Anniversary. Rochester, NY: Eastman Kodak Co., 1988. 5. Feigenbaum H. 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Ultrasonic flow detection: a useful technique in the evaluation of peripheral vascular disease. Am J Surg 1967;113:311–320. 37. Peronneau EPA, Deloche A, Bui-Meng-Hung, et al. Debitmetre ultrasonore: developpmements et applications experimentale. Eur Surg Res 1969;1:147. 38. Kalmanson D, Veyrat C, Derai C, et al. Non-invasive technique for diagnosing atrial septal defect and assessing shunt volume using directional Doppler ultrasound: correlations with phasic flow velocity patterns of the shunt. Br Heart J 1972;34:981–991. 39. Holen J, Simonsen S. Determination of pressure gradient in mitral stenosis with Doppler echocardiography. Br Heart J 1979;41: 529–535. 40. Hatle L, Angelsen B, Tromsdal A. Noninvasive assessment of aortic stenosis by Doppler ultrasound. Br Heart J 1979;43:284–292. 41. Gramiak R, Shah PM, Kramer DH. Ultrasound cardiography: contrast studies in anatomy and function. Radiology 1969;92:939–938. 42. Feigenbaum H, Stone JM, Lee DA, et al. Identification of ultrasound echoes from the left ventricle using intracardiac injections of indocyanine green. Circulation 1979;41:615–621. 43. Seward JB, Tajik AJ, Spangler JG, et al. Echocardiographic contrast studies: initial experience. Mayo Clin Proc 1975;50:163–192. 44. Feinstein SB, Cheirif J, Ten Cate FJ, et al. Safety and efficacy of new transpulmonary ultrasound contrast agent: initial multicenter clinical results. J Am Coll Cardiol 1990;16:316–324. 45. Howry DH, Holmes JH, Cushman RR, et al. Ultrasonic visualization of living organs and tissues, with observations on some disease processes. Geriatrics 1955;10:123. 46. Ebina T, Oka S, Tanaka M, et al. The ultrasono-tomography of the heart and great vessels in living human subjects by means of the ultrasonic reflection technique. Jpn Heart J 1967;8:331–353. 47. Gramiak R, Waag R, Simon W. Cine ultrasound cardiography. Radiology 1973;107:175–180. 48. King DL. Cardiac ultrasonography. Radiology 1972;103:837.B 49. Bom N, Lancee CT, Honkoop J, et al. Ultrasonic viewer for cross-sectional analyses of moving cardiac structures. Biomed Eng 1971;6:500. 50. Griffith JM, Henry WL. A sector scanner for real-time two- dimensional echocardiography. Circulation 1974;49:1147–1152. 51. Eggleton RC, Feigenbaum H, Johnston KW, et al. Visualization of cardiac dynamics with real-time B-mode ultrasonic scanner. In: White D, ed. Ultrasound in Medicine. New York: Plenum Publishing, 1975. 52. Thurstone FL, vonRamm OT. A new ultrasound imaging technique employing two dimensional electronic beam steering. In: Green PS, ed. Acoustical Holography. New York: Plenum Publishing, 1974:149–159. 53. Brandestini MA, Eyer MK, Stevenson JG. M/Q-mode echocardiography: the synthesis of conventional echo with digital multigate Doppler. In: Lancee CT, ed. Echocardiography. The Hague, Netherlands: Martinus-Nijhoff, 1979. 54. Kasai C, Namekawa K, Koyano A, et al. Real-time two- dimensional blood flow imaging using an autocorrelation technique. IEEE Trans Sonics Ultrason 1985;32:460–463. 55. Omoto R, Yokote Y, Takamoto S, et al. The development of real-time two-dimensional Doppler echocardiography and its clinical significance in acquired valvular regurgitation. Jpn Heart J 1984;25:325–340. 56. Frazin L, Talano JV, Stephanides L, et al. Esophageal echocardiography. Circulation 1976;54:102–108. 57. Hisanaga K, Hisanaga A, Nagata K, et al. Transesophageal cross-sectional echocardiography. Am Heart J 1980;100:605–609. 58. Schluter M, Henrath P. Transesophageal echocardiography: potential advantages and initial clinical results. Pract Cardiol 1983;9:149. 59. Souquet J, Hanrath P, Zitelli L, et al. Transesophageal phased array for imaging the heart. IEEE Trans Biomed Eng 1982;29:707. 60. Bom N, Lancee CT, Egmond van FC. An ultrasonic intracardiac scanner. Ultrasonics 1972;10:72–76. 61. Nissen SE, Gurley JC, Grines CL, et al. Intravascular ultrasound assessment of lumen size and wall morphology in normal subjects and patients with coronary artery disease. Circulation 1991;84: 1087. 62. Wollschlager H. Transesophageal echo computer tomography: a new method for dynamic three-dimensional imaging of the heart. In: Computers in Cardiology 1989. IEEE Computer Society, 1990:39. 63. Roelandt J, ten Cate FJ, Bruining N, et al. Transesophageal rotoplane echo-CT. A novel approach to dynamic three-dimensional echocardiography. Thoraxcentre J 1993;6:4–8. 64. vonRamm OT, Smith SW, Pavy HG Jr. High-speed ultrasound volumetric imaging system. Part II. Parallel processing and image display. IEEE Trans Ultrason Ferroelectr Freq Control 1991;38: 109–115. 65. Roelandt J, Wladiniroff JW, Baars AM. Ultrasonic real-time imaging with a hand-held scanner. Ultrasound Med Biol 1978;4:93. Editors: Feigenbaum, Harvey; Armstrong, William F.; Ryan, Thomas Title: Feigenbaum's Echocardiography, 6th Edition Copyright ©2005 Lippincott Williams & Wilkins > Table of Contents > Chapter 2 - Physics and Instrumentation Chapter 2 Physics and Instrumentation Sound is a mechanical vibration transmitted through an elastic medium. When it propagates through the air at the appropriate frequency, sound may produce the sensation of hearing. Ultrasound includes that portion of the sound spectrum having a frequency greater than 20,000 cycles per second (20 KHz), which is considerably above the audible range. The use of ultrasound to study the structure and function of the heart and great vessels defines the field of echocardiography. The production of ultrasound for diagnostic purposes involves physical principles and instrumentation that are both complex and sophisticated. As technology has evolved, a thorough understanding of these principles mandates an extensive background in physics and engineering. Fortunately, the use of echocardiography for clinical purposes does not require a complete mastery of the physics and instrumentation involved in the creation of the ultrasound image. However, a basic understanding of these facts is necessary to take full advantage of the technique and to appreciate the strengths and limitations of the technology. This book is intended principally as a clinical guide to the broad field of echocardiography, to be used by clinicians, students, and sonographers concerned more about the practical application of the technology than the underlying physics. For this reason, an extensive description of the physics and engineering of ultrasound is beyond the scope of this book. Instead, this chapter focuses on those aspects of physics and instrumentation that are relevant to the understanding of ultrasound and its practical application to patient care. In addition, many of the newer technical advances in ultrasound instrumentation are presented briefly, primarily to provide the reader a sense of the changing and ever-improving nature of echocardiography. PHYSICAL PRINCIPLES Ultrasound (in contrast to lower, i.e., audible, frequency sound) has several characteristics that contribute to its diagnostic utility. First, ultrasound can be directed as a beam and focused. Second, as ultrasound passes through a medium, it obeys the laws of reflection and refraction. Finally, targets of relatively small size reflect ultrasound and can therefore be detected and characterized. A major disadvantage of ultrasound is that it is poorly transmitted through a gaseous medium and attenuation occurs rapidly, especially at higher frequencies. As a wave of ultrasound propagates through a medium, the particles of the medium vibrate parallel to the line of propagation, producing longitudinal waves. Thus, a sound wave is characterized by areas of more densely packed particles within the medium (an area of compression) alternating with regions of less densely packed particles (an area of rarefaction). The amount of reflection, refraction, and attenuation depends on the acoustic properties of the various media through which an ultrasound beam passes. Tissues composed of solid material interfaced with gas will reflect most of the ultrasound energy, resulting in poor penetration. Very dense media also reflect a high percentage of the ultrasound energy. Soft tissues and blood allow relatively more ultrasound energy to be propagated, thereby increasing penetration and improving diagnostic utility. Bone also reflects most ultrasound energy, not because it is dense but because it contains so many interfaces. The ultrasound wave is often graphically depicted as a sine wave in which the peaks and troughs represent the areas of compression and rarefaction, respectively (Fig. 2.1 ). Microscopic pressure changes occur within the medium, corresponding to these areas, and result in tiny oscillations of particles, although no actual particle motion occurs. Depicting ultrasound in the form of a sine wave has some limitations but allows the demonstration of several fundamental principles. The sum of one compression and one rarefaction represents one cycle , and the distance between two similar points along the wave corresponds to wavelength (see Table 2.1 for definitions of commonly used terms). Over the range of diagnostic ultrasound, wavelength varies from approximately 0.15 to 1.5 mm in soft tissue. The frequency of the sound wave is P.12 the number of wavelengths per unit of time. Thus, wavelength and frequency are inversely related and their product represents the velocity of the sound wave: FIGURE 2.1. This schematic illustrates how sound can be depicted as a sine wave whose peaks and troughs correspond to areas of compression and rarefaction, respectively. As sound energy propagates through tissue, the wave has a fixed wavelength that is determined by the frequency and amplitude that is a measure of the magnitude of pressure changes. See text for further details. TABLE 2.1 Definitions of Basic Terms Term Definition Absorption The transfer of ultrasound energy to the tissue during propagation Acoustic The product of the density of the medium impedance and the velocity of sound; differences in AI between 2 media determine the ratio of transmitted versus reflected sound at the interface Amplitude The magnitude of the pressure changes along the wave; also, the strength of the wave (in decibels) Attenuation The net loss of ultrasound energy as a wave propagates through a medium Cycle The combination or sum of 1 compression and 1 rarefaction of a propagating wave Dead time The time in between pulses that the echograph is not emitting ultrasound Decibel A logarithmic measure of the intensity of sound, expressed as a ratio to a reference value (dB) Duty factor The fraction of time that the transducer is emitting ultrasound, a unitless number between 0 and 1 Far field The diverging conical portion of the beam beyond the near field Frequency The number of cycles per second, measured in Hertz (Hz) Half-layer value The distance an ultrasound beam penetrates into a medium before its intensity has attenuated to one-half the original value Intensity The concentration or distribution of power within an area, often the cross-sectional area of the ultrasound beam, analogous to loudness Longitudinal A cyclic disturbance in which the energy wave propagation is parallel to the direction of particle motion Near field The proximal cylindrical-shaped portion of the ultrasound beam before divergence begins to occur Period The time required to complete 1 cycle, usually expressed in microseconds (µsec) Piezoelectricity The phenomenon of changing shape in  response to an applied electric current, resulting in vibration and the production of sound waves; the ability to produce an electric impulse in response to a mechanical deformation; thus, the interconversion of electrical and sound energy Power The rate of transfer over time of the acoustic energy from the propagating wave to the medium, measured in Watts Pulse A burst or packet of emitted ultrasound of finite duration, containing a fixed number of cycles traveling together Pulse length The physical length or distance that a pulse occupies in space, usually expressed in millimeters (mm) Pulse repetition The rate at which pulses are emitted from frequency the transducer, ie, the number of pulses emitted within a period of time, usually 1 second Resolution The smallest distance between 2 points that allows the points to be distinguished as separate Sensitivity The ability of the system to image small targets at a given depth Ultrasound A mechanical vibration in a physical medium, characterized by a frequency > 20,000 Hz Velocity The speed at which sound moves through a given medium Wavelength The length of a single cycle of the ultrasound wave; a measure of distance, not time where v is velocity, f is frequency (in cycles per second or hertz) and λ is wavelength. Velocity through a given medium depends on the density and elastic properties or stiffness of that medium. Velocity is directly related to stiffness and inversely related to density. Ultrasound travels faster through a stiff medium, such as bone. Velocity also varies with temperature, but because body temperature is maintained within a relatively narrow range, this phenomenon is of little significance in medical imaging. Table 2.2 provides a comparison of average velocity values in various types of tissues. Within soft tissue, velocity of sound is fairly constant at approximately 1,540 m/sec (or 1.54 m/msec, or 1.54 mm/µsec). Thus, to find the wavelength P.13 of a 3.0-MHz transducer, the solution would be given by: TABLE 2.2 Velocity of Sound in Air and Various Types of Tissues Medium Velocity (m/sec) Air 330 Fat 1450 Water 1480 Soft tissue 1540 Kidney 1560 Blood 1570 Muscle 1580 Bone 4080 A simpler version of this equation is given by λ (in millimeters) = 1.54/f , where f is the transducer frequency (in megahertz). This converts 1,540 m/sec to 1.54 mm/µsec, expresses frequency in megahertz, and yields wavelength in millimeters. Thus, If an ultrasound wave encounters an area of higher stiffness, for example, velocity will increase. Because frequency does not change, wavelength will also increase. As is discussed later, wavelength is a determinant of resolution: the shorter the wavelength is, the smaller the target that is able to reflect the ultrasound wave and thus the greater the resolution. Another fundamental property of sound is amplitude , which is a measure of the strength of the sound wave (Fig. 2.1 ). It is defined as the difference between the peak pressure within the medium and the average value, depicted as the height of the sine wave above and below the baseline. Amplitude is measured in decibels , a logarithmic unit that relates acoustic pressure to some reference value. The primary advantage of using a logarithmic scale to display amplitude is that a very wide range of values can be accommodated and weak signals can be displayed along side much stronger signals. Of practical use, an increase of 6 dB is equal to a doubling of signal amplitude, and 60 dB represents a 1,000-fold change in amplitude or loudness. A parameter closely related to amplitude is power , which is defined as the rate of energy transfer to the medium, measured in watts. For clinical purposes, power is usually represented over a given area (often the beam area) and referred to as intensity (watts per centimeter squared or W/cm2 ). This is analogous to loudness. Intensity diminishes rapidly with propagation distance and has important implications with respect to the biologic effects of ultrasound, which are discussed later. INTERACTION BETWEEN ULTRASOUND AND TISSUE These basic characteristics of ultrasound have practical implications for the interaction between ultrasound and tissue. For example, the higher the frequency of the ultrasound wave (and the shorter the wavelength), the smaller the structures that can be accurately resolved. Because precise identification of small structures is a goal of imaging, the use of high frequencies would seem desirable. However, higher frequency ultrasound has less penetration compared with lower frequency ultrasound. The loss of ultrasound as it propagates through a medium is referred to as attenuation. This is a measure of the rate at which the intensity of the ultrasound beam diminishes as it penetrates the tissue. Attenuation has three components: absorption, scattering, and reflection. Attenuation always increases with depth and is also affected by the frequency of the transmitted beam and the type of tissue through which the ultrasound passes. The higher the frequency is, the more rapidly it will attenuate. Attenuation may be expressed as the “half-value layer†or the “half-power distance,†which is a measure of the distance that ultrasound travels before its amplitude is attenuated to one half its original value. Representative half-power distances are listed in Table 2.3. As a rule of thumb, the attenuation of ultrasound in tissue is between 0.5 and 1.0 dB/cm/MHz. This approximation describes the expected loss of energy (in decibels) that would occur over the round-trip distance that a beam would travel after being emitted by a given transducer. For example, if a 3-MHz transducer is used to image an object at a depth of 12 cm (24-cm round trip), the returning signal could be attenuated as much as 72 dB (or nearly 4,000-fold). As expected, attenuation is greater in soft tissue compared with blood and is even greater in muscle, lung, and bone. The velocity and direction of the ultrasound beam as it passes through a medium are a function of the acoustic impedance of that medium. Acoustic impedance (Z , measured in rayls) is simply the product of velocity (in meters per second) and physical density (in kilograms per cubic meter). As impedance increases, a greater acoustic mismatch is created and relatively more ultrasound energy will be reflected rather than transmitted. Within a homogeneous structure, the density of the medium primarily determines these parameters. In such a structure, sound would travel in a straight line at a constant velocity, depending on the density and stiffness. Within the body, the tissues through which an ultrasound beam passes have different acoustic impedances. When the P.14 beam crosses a boundary between two media, a portion of the energy is reflected, a portion is refracted, and a portion continues in a relatively straight line (Fig. 2.2A ). These interactions between the ultrasound beam with acoustic interfaces form the basis for ultrasound imaging. The phenomena of reflection and refraction obey the laws of optics and depend on the angle of incidence between the transmitted beam and the acoustic interface as well as the acoustic mismatch , i.e., the magnitude of the difference in acoustic impedance. Small differences in velocity also determine refraction. These properties explain the importance of using an acoustic coupling gel during transthoracic imaging. Without the gel, the air-tissue interface at the skin surface results in more than 99% of the ultrasonic energy being reflected at this level. This is primarily due to the very high acoustic impedance of air. The use of gel between the transducer and the skin surface greatly increases the percentage of energy that is transmitted into and out of the body, thereby allowing imaging to occur. TABLE 2.3 Representative Half-Power Distances Relevant to Echocardiography Half-power distance Material (in cm) Water 380 Less attenuation Blood 15 Soft tissue (except 1–5 muscle) Muscle 0.6–1 Bone 0.2–0.7 Air 0.08 Lung 0.05 More attenuation FIGURE 2.2. A : A transmitted wave interacts with an acoustic interface in a predictable way. Some of the ultrasound energy is reflected at the interface and some is transmitted through the  interface. The transmitted portion of the energy is refracted, or bent, depending on the angle of incidence and differences in impedance between the tissues. B : The interaction between an ultrasound wave and its target depends on several factors. A specular reflection occurs when ultrasound encounters a target that is large relative to the transmitted wavelength. The amount of ultrasound energy that is reflected to the transducer by a specular target depends on the angle and the impedance of the tissue. Targets that are small relative to the transmitted wavelength produce a scattering of ultrasound energy, resulting in a small portion of energy being returned to the transducer. This type of interaction results in “speckle†that produces the texture within tissues. As the ultrasound beam is transmitted through tissue, it encounters a complex array of large and small interfaces and targets, each of which affect the transmission of the ultrasound energy. These interactions can be broadly categorized as specular echoes and scattered echoes (Fig. 2.2B ). Specular echoes are produced by reflectors that are large relative to ultrasound wavelength, such as the endocardial surface of the left ventricle. Such targets reflect a relatively greater proportion of the ultrasound energy in an angle-dependent fashion. The spatial orientation and the shape of the reflector determine the angles of specular echoes. Examples of specular reflectors include endocardial and epicardial surfaces, valves, and pericardium. Targets that are small relative to the wavelength of the transmitted ultrasound produce scattering , and such objects are sometimes referred to as Rayleigh scatterers. The resultant echoes are diffracted or bent and scattered in all directions. Because the percentage of energy returning to the transducer from scattered echoes is considerably less than that resulting from specular interactions, the amplitude P.15 of the signals produced by scattered echoes is very low (Fig. 2.2B ). Despite this fact, scattering has important clinical significance (and forms the basis for Doppler imaging). Scattered echoes contribute to the visualization of surfaces that are parallel to the ultrasonic beam and also provide the substrate for visualizing the texture of gray-scale images. The term speckle is used to describe the tissue-ultrasound interactions that result from a large number of small reflectors within a resolution cell. Without the ability to record scattered echoes, the left ventricular wall, for example, would appear as two bright linear structures, the endocardial and the epicardial surfaces, with nothing in between. From the above discussion, it is evident that the interaction between an ultrasound beam and a reflector depends on the relative size of the targets and the wavelength of the beam. If a solid object is submerged in water, for example, whether reflection of ultrasound occurs depends on the size of the object with respect to the wavelength of the transmitted ultrasound. Specifically, the thickness or profile of the object relative to the ultrasound beam must be at least one-fourth the wavelength of the ultrasound. Thus, as the size of the target decreases, the wavelength of the ultrasound must decrease proportionately to produce a reflection and permit the object to be recorded. This explains why higher frequency ultrasound allows smaller objects to be visualized. In clinical practice, echocardiography typically employs ultrasound with a range of 2,000,000 to 8,000,000 cycles per second (2–8 MHz). At a frequency of 2 MHz, it is generally possible to record distinct echoes from interfaces separated by approximately 1 mm. However, because high-frequency ultrasound is reflected by many small interfaces within tissue, resulting in scattering, much of the ultrasonic energy becomes attenuated and less energy is available to penetrate deeper into the body. Thus, penetration is reduced as frequency increases. Similarly, as the medium becomes less homogeneous, the degree of reflection and refraction increases, resulting in less penetration of the ultrasound energy.  FIGURE 2.3. The principles of piezoelectricity are illustrated in this schematic. A piezoelectric crystal will vibrate when an electric current is applied, resulting in the generation and transmission of ultrasound energy. Conversely, when reflected energy encounters a piezoelectric crystal, the crystal will change shape in response to this interaction and produce an electrical impulse. See text for further details. THE TRANSDUCER The use of ultrasound for imaging became practical with the development of piezoelectric transducers. The principles of piezoelectricity are illustrated in Figure 2.3. Piezoelectric substances or crystals rapidly change shape or vibrate when an alternating electric current is applied. It is the rapidly alternating expansion and contraction of the crystal material that produces the sound waves. Equally important is the fact that a piezoelectric crystal will produce P.16 an electric impulse when it is deformed by reflected sound energy. Such piezoelectric crystals form the critical component of ultrasound transducers. Although a variety of piezoelectric materials exist, most commercial transducers employ ceramics, such as ferroelectrics, barium titanate, and lead zirconate titanate. The creation of an ultrasound pulse thus requires that an alternating electric current be applied to a piezoelectric element. This results in the emission of sound energy from the transducer, followed by a period of quiescence during which the transducer “listens†for some of the transmitted ultrasound energy to be reflected back (known as “dead time†). The amount of acoustic energy that returns to the transducer is a measure of the strength and depth of the reflector. The time required for the ultrasound pulse to make the round-trip from transducer to target and back again allows calculation of the distance between the transducer and reflector. An ultrasound transducer consists of many small, carefully arranged piezoelectric elements that are interconnected electronically (Fig. 2.4 ). The frequency of the transducer is determined by the thickness of these elements. Each element is coupled to electrodes, which transmit current to the crystals, and then record the voltage generated by the returning signals. An important component of transducer design is the dampening (or backing) material, which shortens the ringing response of the piezoelectric material after the brief excitation pulse. An excessive ringing response (or “ringdown†) lengthens the ultrasonic pulse and decreases range resolution. Thus, the dampening material both shortens the ringdown and provides absorption of backward and laterally transmitted acoustic energy. At the surface of the transducer, matching layers are applied to provide acoustic impedance matching between the piezoelectric elements and the body. This increases the efficiency of transmitted energy by minimizing the reflection of the ultrasonic wave as it exits the transducer surface. Transducer design is critically important to optimal image creation. An important feature of ultrasound is the ability to direct or focus the beam as it leaves the transducer. This results in a parallel and cylindrically shaped beam. Eventually, however, the beam diverges and becomes cone shaped (Fig. 2.5 ). The proximal or cylindrical portion of the beam is referred to as the near field or Fresnel zone. When it begins to diverge, it is called the far field or Fraunhofer zone. For a variety of reasons, imaging is optimal within the near field. Thus, maximizing the length of the near field is an important goal of echocardiography. FIGURE 2.4. A schematic diagram of a transducer is provided. See text for details. The length of the near field (l ) is described by the formula: where r is the radius of the transducer and λ is the wavelength of the emitted ultrasound. Either decreasing the wavelength (increasing the frequency) or increasing the size of the transducer will lengthen the near field. These relationships are illustrated in Figure 2.6. From the above formula, one might conclude that optimal ultrasound imaging would always employ a large-diameter, high-frequency transducer to maximize the length of the near field. Several factors prevent this approach from being practical. First, the transducer size is predominantly limited by the size of the intercostal spaces. A transducer that is too large will not be able to image between the ribs. Second, although higher frequency does lengthen the near field, it also results in greater attenuation and lower penetration of the ultrasound energy, thereby limiting its usefulness. These tradeoffs must be balanced to maximize imaging performance. Even when the near field length is maximized, most targets will still lie in the far field. To improve imaging in this area, the rate of beam divergence must be minimized. To decrease the amount of divergence in the far field, a large-diameter, high-frequency transducer is optimal. As discussed previously, focusing of P.17 the transmitted beam tends to improve imaging in the near field but will increase the rate or angle of divergence in the far field (Fig. 2.7 ). Focusing is accomplished through the use of an acoustic lens placed on the surface of the transducer or by constructing the piezoelectric crystal in a concave shape. Thus, transducer frequency, size, and focusing all interact to affect image quality in the near and far fields. Tradeoffs exist that must be taken into account to create optimal images. Figure 2.8 is an example of the effects of varying transducer frequencies on image quality and appearance. On the left, a short-axis view is recorded using a 3.0-MHz transducer. On the right, a similar image is captured using a 5.0-MHz probe. Note how the higher frequency results in improved resolution and detail, especially within the myocardium. FIGURE 2.5. When ultrasound is emitted from a transducer, the shape of the beam obeys particular physical principles. If the transducer face is round, the transmitted beam will remain cylindrical for a distance, defined as the near field. After a particular distance of propagation, the beam will begin to diverge and become cone shaped. This region of the beam is referred to as the far field. Within this portion of the beam, a decrease in intensity occurs. The length of the near field is determined by the radius of the transducer face and the wavelength or frequency of the transmitted energy. See text for details. FIGURE 2.6. The length of the near field depends on transducer frequency and transducer size, as illustrated in these four examples. On the left, a transducer with a 10 mm diameter emits ultrasound at 2.0 MHz. This determines both the length of the near field and the rate of divergence in the far field. If the same size transducer emits energy at 4 MHz, the length of the near field increases and the rate of dispersion is less. A transducer half that size (5 mm) transmitting at 4.0 MHz will have a shorter near field. Finally, a 5 mm transducer that transmits at 2 MHz will have the shortest near field and the greatest rate of dispersion in the far field.  FIGURE 2.7. The ultrasound beam emitted by a transducer can be either unfocused (top) or can be focused by use of an acoustic lens (bottom). Focusing results in a narrower beam but does not change the length of the near field. An undesirable effect of focusing is that the rate of dispersion in the far field is greater. MANIPULATING THE ULTRASOUND BEAM For most clinical applications, the ultrasound beam is both focused and steered electronically. Although beam manipulation can be done mechanically, with modern equipment, it is primarily achieved through the use of phased-array transducers, which consist of a series of small piezoelectric elements interconnected electronically (Fig. 2.9 ). In such transducers, the wave front of the beam consists of the sum of the individual wavelets produced by each element. By manipulating the timing of excitation of individual elements, both focusing and steering are possible. If all elements are excited simultaneously, each one will produce a circular wavelet that combines to generate a longitudinal wave front that is parallel to the face of the transducer and propagates in a direction perpendicular to that face. By adjusting the timing of excitation, as shown in Figure 2.10A , the beam can be steered. Further adjustments in the timing allow the beam to be steered through a sector arc, resulting in a two-dimensional image. Using a similar approach, electronic transmit focusing of the beam is also possible (Fig. 2.10B ). For example, by exciting the outside elements first and then progressively activating the more central elements, the individual wavelets form a curved front that allow focusing at a particular distance within the near field. This can either be fixed or adjustable, and the process is referred to as dynamic transmit focusing. It should be recognized that the ultrasound beam is a three- dimensional structure that, in the case of a phased-array transducer, is roughly rectangular in cross section (Fig. 2.11 ). The dimensions of the beam are referred to as axial (along the axis of wave propagation) and lateral (parallel to the face of the transducer, sometimes called azimuthal). The lateral dimension is further divided into a vertical and horizontal component. Acoustic focusing through a lens will change the shape in the vertical and horizontal dimensions equally. Electronic focusing will narrow the beam in one of these two dimensions, resulting in a “thinner†sector slice. Transducers that employ anular phased-array technology have the capacity to focus in both dimensions, resulting in a compact, high-intensity beam profile. Another type of transducer uses a linear array of elements. Such transducers have a rectangular face with P.18 crystals aligned parallel to one another along the length of the transducer face. Unlike phased-array transducers, the elements are excited simultaneously so the individual scan lines are directed perpendicular to the face and remain parallel to each other. This results in a rectangle-shaped beam that is unfocused. Linear-array technology is often used for abdominal, vascular, or obstetric applications. Alternatively, the face of a linear transducer can be curved to create a sector scan. This innovative design is now being used in some handheld ultrasound devices. FIGURE 2.8. The effects of different transducer frequencies on image quality and appearance are demonstrated. A : A 3.0-MHz  transducer is used to record a short-axis view. B : The same image is recorded using a 5.0-MHz transducer. FIGURE 2.9. A phased-array ultrasound transducer is shown. Focusing has the effect of concentrating the acoustic energy into a smaller area, resulting in increased intensity at the point of focus. Intensity also varies across the lateral dimensions of the beam, being greatest at the center and decreasing in intensity toward the edges. When the shape of the ultrasonic beam is diagrammed, it is conventional P.19 to draw the edge of the beam to the half-value limit of the beam plot. An example of a transaxial beam plot is illustrated in Figure 2.12. This diagram illustrates the important relationship between intensity and beam width. At its peak intensity, the beam may be as narrow as 1 mm. At its weakest intensity, however, beam width may be as great as 12 mm. For purposes of comparison, it is customary to measure the beam width at its half amplitude or intensity. In the example shown, the beam width would be reported as 6.2 mm. Finally, it should be remembered that gain setting will affect these values in a predictable manner. At high gain settings, the weaker portion of the ultrasound beam is recorded and beam width is greater. Conversely, at low gain settings, the beam width would be narrower. FIGURE 2.10. A : Phased-array technology permits steering of the ultrasound beam. By adjusting the timing of excitation of the individual piezoelectric crystals, the wave front of ultrasound energy can be directed, as shown. Beam steering is a fundamental feature of how two-dimensional images are created. B : By adjusting the timing of excitation of the individual crystals within a phased-array transducer, the beam can be focused. In this example, the outer elements are fired first, followed sequentially by the more central elements. Because the speed of sound is fixed, this manipulation in the timing of excitation results in a wave front that is curved and focused. This is called transmit focusing.  FIGURE 2.11. The ultrasound beam can be represented as a three-dimensional structure. A single-crystal transducer (top) will emit a cylindrically shaped beam. If the transducer face is rectangular shaped (bottom) , the beam will also have a rectangular shape. The various beam axes are labeled in the two drawings. As is apparent from the previous discussion, focusing of the ultrasonic beam is generally desirable. By increasing beam intensity within the near field, the strength of returning signals is enhanced. An undesirable effect of focusing is its effect on beam divergence in the far field. Because focusing results in a beam with a smaller radius, the angle of divergence in the far field is increased. However, because beam divergence begins from a small cross-sectional area of a focused beam, the net effect is variable. The result of these relationships is a tradeoff between resolution at the point of focus and depth of field. Divergence also contributes to the formation of important imaging artifacts such as side lobes (discussed later).  FIGURE 2.12. A transaxial beam plot is demonstrated. The beam width or lateral resolution is a function of the intensity of the ultrasonic beam. The beam width is commonly measured at the half-intensity level, and, in this case, the beam width would be reported as 6.2 mm. RESOLUTION Resolution is the ability to distinguish between two objects in close proximity. Because echocardiography depends on its ability to image small structures and provide detailed anatomic information, resolution is one of its most important variables. Furthermore, because echocardiography is a dynamic imaging technique, resolution has at least two components: spatial and temporal. Spatial resolution is defined as the smallest distance that two targets can be separated for the system to distinguish between them. It, too, has two components: Axial resolution refers to the ability to differentiate two structures lying along the axis of the ultrasound beam (i.e., one behind the other) and lateral resolution refers to the ability to distinguish two reflectors that lie side by side relative to the beam (Fig. 2.13 ). The primary determinants of axial resolution are the frequency of the transmitted wave and, more importantly, its effect on pulse length. Higher frequency is associated with shorter wavelength, and the size of the wave relative to the size of the object determines resolution. In addition to frequency, pulse length or duration also affects axial resolution. The shorter the train of cycles is, the greater the likelihood that two closely positioned targets can be resolved. Because a higher frequency and/or broad bandwidth transducer delivers a shorter pulse, it is also associated with higher resolution. FIGURE 2.13. The different types of resolution are demonstrated in this schematic. See text for details. P.20 Lateral resolution varies throughout the field of imaging and is affected by several factors. The width or thickness of the interrogating beam, at a given depth, is the most important determinant. Ideally, the ultrasonic beam should be very narrow to provide a thin “slice†of the heart. Recall that the beam has finite width, even in the near field, and tends to diverge as it propagates. The importance of beam width stems from the fact that the system will display all targets within the path of the beam along a single line represented by the central axis of the beam. In other words, the echograph displays structures within the image as if the beam were infinitely narrow. Thus, lateral resolution diminishes as beam width (and depth) increases. The distribution of intensity across the beam profile will also affect lateral resolution. As illustrated in Figure 2.14 , both strong and weak reflectors can be resolved within the central portion of the beam, where intensity is greatest. At the edge of the beam, however, only relatively strong reflectors may produce a signal. Furthermore, the true size and position of such objects may be distorted by the width of the beam, resulting in significant beam width artifacts. This is illustrated in Figure 2.15. This observation also explains the importance of overall system gain and its effect on lateral resolution. Gain is the amplitude, or the degree of amplification, of the received signal. When gain is low, weaker echoes from the edge of the beam may not be recorded and the beam appears relatively narrow. If system gain is increased, weaker and more peripheral targets are recorded and beam width appears greater. Thus, to enhance lateral resolution, a minimal amount of system gain should be employed. Figure 2.16 illustrates how changes in gain setting can drastically alter lateral resolution and anatomic information. FIGURE 2.14. This schematic illustrates the interrelationship between beam intensity and acoustic impedance. The center of the beam has higher intensity compared with the edges. A : Whether an echo is produced, and with what amplitude it is recorded, depends on the relationship between intensity and acoustic impedance. Objects with higher impedance (black dots) produce stronger echoes and can therefore be detected even at the edges of the beam. Weaker echo-producing targets (gray dots) only produce echoes when they are located in the center of the beam. B : The effect of beam width on target location is shown. Objects A and B are nearly side by side with B slightly farther from the transducer. Because of the width of the beam, both objects are recorded simultaneously. The resulting echoes suggest that the two objects are directly behind each other (A' and B') rather than side by side. A third component of resolution is called contrast resolution. Contrast resolution refers to the ability to distinguish and to display different shades of gray within the image. This is important both for the accurate identification of borders and for the ability to display texture or detail within the tissues. To convert the returning radio frequency (RF) information into a gray-scale image, pre- and post-processing of the data are performed. These steps in image formation rely heavily on contrast resolution. From a practical standpoint, contrast resolution is necessary to differentiate tissue signals from background noise. Contrast resolution is also dependent on target size. A higher degree of contrast is needed to detect small structures compared with larger targets. Temporal resolution , or frame rate, refers to the ability of the system to accurately track moving targets over P.21 time. It is dependent on the amount of time required to complete a scan, which in turn is related to the speed of ultrasound and the depth of the image as well as the number of lines of information within the image. Generally, the greater the number of frames per unit of time, the smoother and more aesthetically pleasing the real-time image. Factors that reduce frame rate, such as increasing depth of field, will diminish temporal resolution. This is particularly important for structures with relatively high velocity, such as valves. Temporal resolution is the main reason that M-mode echocardiography is still a useful clinical tool. With sampling rates of 1,000 to 2,000 images per second, temporal resolution of this modality is much higher than that of two-dimensional imaging. FIGURE 2.15. These diagrams demonstrate how beam width distorts the image in a two-dimensional sector scan (A) and how a focused beam can reduce this distortion (B). The true image should be a series of dots; beam width, however, distorts the image into a series of dashes. CREATING THE IMAGE The instrument used to create an ultrasound image is called an echograph. It contains the electronics and circuitry needed to transmit, receive, amplify, filter, process, and display the ultrasound information. The essential P.22 components of the system are illustrated in Figure 2.17. As a first step, the returning energy is converted from sound waves to voltage signals. These are very low amplitude, high-frequency signals that must be amplified and, because they arrive slightly out of phase, realigned in time. In modern instrumentation, this realignment is accomplished using a digital beam former to allow proper summation and phasing of all the channels. Because the signals are still very high frequency at this point, the scan lines are referred to as RF data. The complexity of the information at this stage is in part due to the wide range of amplitudes and the inclusion of background noise. Logarithmic compression and filtering are performed to render the RF data more suitable for processing. FIGURE 2.16. Parasternal long-axis images demonstrate the effect of gain on the appearance on the echocardiographic image. A : Gain is adjusted appropriately to allow recording of all relevant information. B : Too much gain is used, distorting the image, reducing resolution, and increasing noise. The polar scan line data at this point consist of sinusoidal waves, and each ultrasound target is represented as a group of these high- frequency spikes. Each group of high-frequency RF data is consolidated into a single envelope through a curve-fitting process called envelope detection. The resulting signal is then referred to as the polar video signal. This is sometimes called R-theta, indicating that each point in a polar map can be defined by its distance (R) and angle (theta) from a reference point. The next very important step involves digital scan conversion and refers to the complex task of converting polar video data into a cartesian or rectangular format. The image formed at this stage can be either stored in digital format or converted to analog data for videotape storage and display. Figure 2.18 displays these different forms of imaging data as energy is received and processed by the echograph. The energy created by excitation of the piezoelectric elements is an RF signal (Fig. 2.18A ). As discussed in the previous section, for the signal to be in a form that can be displayed visually, it must be converted to a video signal. This is accomplished by outlining (envelope detection) the outer edge of the upper portion, or positive deflection, of the RF signal (Fig. 2.18B ). Differentiation of the video signal effectively accentuates the leading edge of the echo (Fig. 2.18C ), providing a brighter signal and improving the ability to differentiate closely spaced targets. This is sometimes referred to as A-mode, for amplitude, P.23 imaging. Finally, intensity modulation converts the height or amplitude of the signal to a corresponding brightness level for video display (Fig. 2.18D ). This is often called B-mode, for brightness, imaging and forms the basis of both M-mode and two-dimensional imaging display. How these various signal formats are used to create a visual display is covered in greater detail in a later section. FIGURE 2.17. A block diagram shows the components of an echograph. The various steps needed to create an image, beginning at the transducer and continuing to the display, are included. See text for details.  FIGURE 2.18. Some of the key steps in image creation are illustrated graphically. See text for details. TRANSMITTING ULTRASOUND ENERGY For most clinical applications, ultrasound is emitted from the transducer as a brief pulse of energy. A fundamental control feature is power output, which is simply the amount of ultrasound energy within each emitted pulse. In general, the higher the power output, the higher the amplitude of the returning signal. The pulse, which is a collection of cycles traveling together, is emitted at fixed intervals (Fig. 2.19 ). The time between pulsing is referred to as the dead time and is largely a function of depth. During the dead time, the transducer is “listening†for returning signals. The duration of the ultrasound pulse is sometimes referred to as pulse length , and the pulse repetition period represents the total of one pulse length plus one dead time. To image at a greater depth, the dead time is lengthened, allowing the ultrasound system to listen for reflections arising from greater depths before returning to the transducer. Duty factor , or the percentage of time that the transducer is pulsing, is simply the pulse duration divided by the pulse repetition period. This is a very small number, in the range of 0.1%, indicating that the system is “on†for a brief time and “off,†or listening, for the majority of time.