CT Imaging : Practical Physics, Artifacts, and Pitfalls 9780199987993, 9780199782604

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CT IMAGING

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CT IMAGING PRACTICAL PHYSICS, ARTIFACTS, AND PITFALLS Editor: Alexander C. Mamourian MD Professor of Radiology Division of Neuroradiology Department of Radiology Perelman School of Medicine of the University of Pennsylvania Philadelphia, Pennsylvania

Contributors: Harold Litt MD, PhD Nicholas Papanicolaou MD, FACR Assoc. Professor of Radiology and Medicine Chief, Cardiovascular Imaging Department of Radiology Perelman School of Medicine of the University of Pennsylvania Philadelphia, Pennsylvania

Co-Chief, Body CT Section Professor of Radiology Department of Radiology Perelman School of Medicine of the University of Pennsylvania Philadelphia, Pennsylvania

Supratik Moulik MD

Josef P. Debbins PhD, PE, DABMP

Fellow, Cardiovascular Imaging Department of Radiology Perelman School of Medicine of the University of Pennsylvania Philadelphia, Pennsylvania

Staff Scientist Keller Center for Imaging Innovation Department of Radiology St. Joseph’s Hospital and Medical Center Phoenix, Arizona

1 2013

1

Oxford University Press is a department of the University of Oxford. It furthers the University’s objective of excellence in research, scholarship, and education by publishing worldwide Oxford New York Auckland Cape Town Dar es Salaam Hong Kong Karachi Kuala Lumpur Madrid Melbourne Mexico City Nairobi New Delhi Shanghai Taipei Toronto With offices in Argentina Austria Brazil Chile Czech Republic France Greece Guatemala Hungary Italy Japan Poland Portugal Singapore South Korea Switzerland Thailand Turkey Ukraine Vietnam

© Oxford University Press 2013 Published in the United States of America by Oxford University Press 198 Madison Avenue, New York, New York 10016 www.oup.com Oxford is a registered trademark of Oxford University Press in the UK and certain other countries All rights reserved. No part of this publication may be reproduced, stored in a retrieval system, or transmitted, in any form or by any means, electronic, mechanical, photocopying, recording, or otherwise, without the prior permission of Oxford University Press. Library of Congress Cataloging-in-Publication Data CT imaging : practical physics, artifacts, and pitfalls / editor, Alexander C. Mamourian; contributors, Harold Litt ... [et al.]. p. ; cm. Includes bibliographical references and index. ISBN 978-0-19-978260-4(pbk. : alk. paper) I. Mamourian, Alexander C. II. Litt, Harold I. [DNLM: 1. Tomography, X-Ray Computed. 2. Cardiac Imaging Techniques. 3. Nervous System—radiography. 4. Radiation Dosage. 5. Radiation Protection. 6. Whole Body Imaging. WN 206] LC Classification not assigned 616.07'5722—dc23 2012038160

1 3 5 7 9 8 6 4 2 Printed in the United States of America on acid-free paper

CONTENTS Introduction Acknowledgements Dedication

vii ix xi

1

HISTORY AND PHYSICS OF CT IMAGING Alexander C. Mamourian

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RADIATION SAFETY AND RISKS Alexander C. Mamourian and Josef P. Debbins

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CARDIAC CT IMAGING TECHNIQUES Supratik Moulik and Harold Litt

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CARDIAC CT ARTIFACTS AND PITFALLS Supratik Moulik and Harold Litt

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NEURO CT ARTIFACTS Alexander C. Mamourian

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NEURO CT PITFALLS Alexander C. Mamourian

147

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BODY CT ARTIFACTS Nicholas Papanicolaou

197

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BODY CT PITFALLS Nicholas Papanicolaou

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TEST QUESTIONS Alexander C. Mamourian

225

Index

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INTRODUCTION I could say that computed tomography (CT) and my career started together, since the first units arrived in most hospitals the same year that I entered my radiology residency. But while I knew the physics of CT well at that time, over the next 30 years CT became increasingly complicated in a quiet sort of way. While MR stole the spotlight during much of that time, studies that were formerly unthinkable, like CT imaging the heart and cerebral vasculature, have become routine in clinical practice. But these expanding capabilities of CT have been made possible by increasingly sophisticated hardware and software. And while most manufacturers provide a clever interface for their CT units that may lull some into thinking that things are under control, the user must understand both the general principles of CT as well as the specific capabilities of their machine because of the potential to harm patients with X-rays. For example, it was reported not long ago that hundreds of patients received an excessive X-ray dose during their CT brain perfusion exams. Although that was troubling enough, the unusually high dose was eventually attributed in some share to the well-meaning but improper use of software commonly used to reduce patient X-ray dose but only for specific applications that do not include perfusion. This book was never intended to be the defi nitive text on the history, physics, and techniques of CT scanning. Our goal was to offer a collection of useful advice taken from our experience about modern CT imaging for an audience of radiology residents, fellows, and technologists. It was an honor and a pleasure to work with my co-authors, an all-star cast of experts in this field, and it is our collective hope you will fi nd this book helpful in the same way that the owner’s manual that comes with a new car is helpful; not enough information to rebuild the engine, but what you need to reset the clock when daylight saving rolls around or change the oil. Many experienced CT users will very likely fi nd some things useful here as well. The review of CT hardware in Chapter 1 should get you off to a good start since the early scanners were just simpler and for that reason easier to understand. The following chapters build on that foundation. Chapter 2 provides a review of the language of X-ray dose and dose reduction, followed by a comprehensive description of the advanced techniques used for cardiac CT in Chapter 3. Feel free at any time to explore the cases in Chapters 4 through 8. Most of these include discussions of practical physics appropriate to that particular artifact or pitfall. In the fi nal chapter, you will fi nd 10 questions that will test your understanding of CT principles. Take it at the start or at the end to see how you stand on this topic. While there is a rationale to the arrangement of the book you may want to keep it nearby and go to appropriate chapters for those questions that may arise about CT dose, protocols, and artifacts in your daily practice. If you get nothing else from reading this book, you should be sure to learn the language of CT dose explained in Chapter 2. Understanding radiation dose specific to CT has become more important than ever in this time of increasing patient awareness, CT utilization, and availability of new software tools for dose reduction. We hope that this book will help you to create the best possible CT images, at the lowest possible dose, for your patients.

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ACKNOWLEDGMENTS I want to thank Cheryl Boghosian and Neil Roth in New Hampshire, for their wonderful hospitality, generous spirit, and faithful friendship over many years, and most recently for giving me the time and space to fi nish this book. My sincere thanks also go to Andrea Seils at Oxford Press. Every writer should be blessed with an editor of her caliber. I will be forever grateful to Dr. Robert Spetzler and all the staff at the Barrow Neurological Institute for giving me the inspiration and the opportunity to write at all.

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DEDICATION I dedicate this book to my parents, Marcus and Maritza, who have given unselfi shly of themselves to so many. To Pamela, Ani, Molly, Elizabeth, and Marcus, I can fi nd no words that can express my endless affection and gratitude.

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HISTORY AND PHYSICS OF CT IMAGING Alexander C. Mamourian

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CT IMAGING

The discovery of X-rays over 100 years ago by Wilhelm Roentgen marks the stunning beginning of the entire field of diagnostic medical imaging. While the impact of his discovery on the fields of physics and chemistry followed, the potential for medical uses of X-rays was so apparent from the start that, within months of his fi rst report, the fi rst clinical image was taken an ocean away in Hanover, New Hampshire. A photograph of that particular event serves as a reminder of how naïve early users of X-ray were with regard to adverse effects of radiation (Figure 1.1). We can only hope that our grandchildren will not look back at our utilization of CT in quite the same way. Although plain X-ray images remain the standard for long bone fractures and preliminary chest examinations, they proved to be of little value for the diagnosis of diseases involving the brain, pelvis, or abdomen. This is because conventional X-ray images represent the net attenuation of all the tissue between the X-ray source and the fi lm (Figures 1.2–1.4). This inability to differentiate tissues of similar density on X-ray is due in part to the requirement for the X-ray beam to be broad enough to cover all the anatomy at once. As a result of this large beam, many of the X-rays that are captured on film have been diverted from their original path into other directions, and these scattered X-rays limit the contrast between similar tissues. This problem was well known to early imagers, and, prior to the invention of computed tomography (CT), a number of solutions were proposed to accentuate tissue contrast on X-ray images. The most effective of these was a device that linked the X-ray tube and film holder together, so that they would swing back and forth in reciprocal directions on either side of patient, around a single pivot point. This was effective to some

Figure 1.1 This photograph captures the spirit of early X-ray exams. Note the pocket watch used to time the exposure (left ) and the absence of any type of radiation protection for the patient or observers. The glowing cathode ray tube (positioned over the arm of the patient, who is sitting with his back to the photographer) was borrowed from the department of physics at Dartmouth College. As rudimentary as this apparatus might appear, it was effective in demonstrating the patient’s wrist fracture. Image provided courtesy of Dr. Peter Spiegel, Dartmouth-Hitchcock Medical Center, Lebanon, New Hampshire.

History and Physics of CT Imaging

(A)

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Film

Net attenuation X-ray 1

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beam

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Figure 1.2 While X-ray images (A) are useful for demonstrating contrast between bone, soft tissue, and air, they are not effective at showing contrast between tissues of similar attenuation. In this image, the pancreas, liver, and kidneys cannot be identified separately because they all blend with nearby tissues of similar density. That is in part because the flat X-ray image can only show the net attenuation of all the tissues between the X-ray source and the film or detector. This is illustrated mathematically in B, where these two rows of blocks of varying attenuation would nevertheless have the same net attenuation on a conventional X-ray image.

degree because it created blurring of the tissues above and below the pivot plane (Figure 1.5), and this technique became know as simply tomography. When I was a resident, we used several variations of this technique for imaging of the kidneys and temporal bones to good effect since the tissues in the plane of the pivot point were in relatively sharp focus, at least sharper than conventional X-rays. Computed tomography proved to be much more than an incremental advance over simple X-ray tomography, however. That is because it both improved tissue contrast and, for the fi rst time, allowed imagers to see the patient in cross-section. The remarkable sensitivity to tissue contrast offered by CT was in some sense serendipitous since it was the byproduct of the use a very narrow X-ray beam for data collection (Figure 1.6). This narrow beam, unlike the wide X-ray beam used for plain fi lms, significantly reduces scatter radiation. For physicians familiar with conventional X-ray images, those early CT images were really just as remarkable as Roentgen’s original X-ray images. The benefits offered by CT imaging to health care was formally acknowledged with the 1979 Nobel Prize for medicine going to Godfrey Hounsfield, just 6 years after his fi rst report of it. The prize was shared with Allan Cormack, in recognition of his contributions to the process of CT image reconstruction. But this prestigious award was not necessary to bring public attention to this new imaging device. At the time the Nobel was awarded, there were already over 1,000 CT units operating or on order worldwide. At the time of his discovery, Godfrey Hounsfield was employed by a British fi rm called EMI (Electrical and Musical Industries) that had interests in both music and musical hardware. While EMI is better known now for its association with both Elvis Presley and the Beatles, it was much more than a small recording company with some good fortune in signing future stars. EMI manufactured a broad range of electrical hardware, from record players to giant radio transmitters, and

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Figure 1.3 and 1.4 Another significant limitation of plain film is that there is no indication of depth even when sufficient image contrast is present. For example, on this single plain film of the skull it appears at first glance that this patient’s head is full of metal pins (1.3). This is because an X-ray image is just a two-dimensional representation of a three-dimensional object, and each point on the image reflects the sum attenuation of everything that lies between the X-ray source and that point on the film. While you can easily see that there are a large number of metal pins superimposed on the skull in this example, you cannot tell whether they are on top of the skull, behind the skull, or inside the skull (perhaps from some terrible industrial accident). The computed tomography (CT) image of this patient shows that they are, fortunately, hairpins that are outside the skull (Figure 1.4; arrows).

a fortuitous and unusual combination of broad interests in electronics with substantial financial support offered by its music contracts apparently gave Hounsfield the latitude necessary for his distinctly unmusical research into CT imaging. In his lab, he built a device intended to measure the variations in attenuation across a phantom using a single gamma ray source and single detector. Gamma rays are, of course, naturally occurring radiation, and so the fi rst device he built did not use an X-ray tube at all but a constrained radioactive element. By measuring precisely how much the phantom attenuated the gamma rays in discrete steps from side to side, and then repeating those measurements in small degrees of rotation around the object, Hounsfield showed that it was possible to recreate the internal composition of a solid phantom using exclusively external measurements. While CT is commonplace now, at the start this capability to see inside opaque objects must have seemed analogous to Superman’s power to see through solid walls. That large dataset collected by Hounsfield’s device was then converted into an image using known mathematical calculations (Figures 1.7, 1.8) with the aid of a computer of that era. Computed tomography was initially considered to be a variation of existing tomography, so it was called “computed” tomography, or more accurately computed axial tomography aka CAT scanning. This acronym was commonly a source of humor when confused with the pet (no pun intended), and eventually it was shortened to just “CT.” Hounsfield was honored for the creation of this remarkable imaging tool by having the standard unit of CT attenuation named a “Hounsfield unit,” which is abbreviated HU.

History and Physics of CT Imaging

Figure 1.5 This drawing from a patent illustration shows the complex mechanics of a tomography device. In this design, the X-ray tube is under the patient table and the film above. The belt at the bottom drives the to-and-fro movement of the entire apparatus. From AG Filler. The history, development and impact of computed imaging in neurological diagnosis and neurosurgery: CT, MRI, and DTI. Doi:10.103/npre.2009.3267.5

The medical implications of his device were quite evident to Hounsfield from his earliest experiments, and EMI was supportive of his research in this direction. As the invention moved into a clinical imaging tool, the mathematical reconstruction used for initial experiments proved to be too time-consuming using the computers available at that time. Faster reconstruction was essential for clinical use and, in recognition of his research that contributed to the faster reconstruction speeds for CT, Allan Cormack was also recognized with a share of the 1979 Nobel Prize. In common with many scientific advances, Cormack’s investigations preceded the invention of CT imaging by many years. It was twenty years prior to Hounsfield’s work, after the resignation of the only other nuclear physicist in Capetown, South Africa, that Cormack became responsible for the supervision of the radiation therapy program at a nearby hospital. Without a dedicated medical background, he brought a fresh perspective on his new responsibilities and was puzzled at the usual therapy planning process used at that time. It presumed that the human body was homogeneous as far as X-rays are concerned, when it clearly was not. He thought that if the tissue-specific X-ray attenuation values for different tissues were known, it would eventually be of benefit not only for therapy but also for diagnosis. He eventually published his work on this subject in 1963, nearly a decade prior to Hounsfield’s first report of his CT device. In his Nobel acceptance lecture, Cormack reflected that, immediately after the publication of his work, it received little attention except from a Swiss center for avalanche prediction that hoped it would prove to be of value for their purposes. It did not.

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Figure 1.6 This early CT of the brain allowed the imager to see the low attenuation CSF within the ventricles as well as the high attenuation calcifications in the ventricular wall in this patient with tuberous sclerosis.

Axial Versus Helical Imaging While early CT scanners were quite remarkable in their time, they were really quite slow as they went about their businesslike “translate-rotate” method of data collection. For example, it took about 5 minutes to accumulate the data for two thick (>10mm) slices of the brain at an 80 ×80 matrix. While still remarkable at that time, these scanners were deemed inadequate for much else apart from brain imaging. Even with their limitations, early EMI CT scanners were very expensive, costing about $300,000 dollars even in 1978, and that got the attention of many other manufacturers around the world. It became a race among them to establish a foothold in this lucrative new market. As a result of this concerted effort, CT scan times dropped rapidly as manufacturers offered faster and better units; as a result, it was not long before EMI was left behind. Those fi rst-generation scanners were made obsolete by faster “second-generation” units that used multiple X-ray sources and detectors. Not long afterward, these second-generation scanners were surpassed by scanners using what we call “third-generation” design, which eliminated the “translate” movement. Now the X-ray fan beam, along with its curved detector row (Figure 1.9), could spin around the patient without stopping. That design still remains the preferred arrangement on current scanners since it readily accommodates large X-ray tubes, both axial and helical imaging, and wide detector arrays. Since they spin together, the large detector arrays nicely balance the large X-ray tubes.

History and Physics of CT Imaging

Figure 1.7 Hounsfield’s patent on CT included an illustration (upper left drawing labeled A) of the lines of data that were collected in a translate-rotate pattern, shown here for only three different angles. From AG Filler. The history, development and impact of computed imaging in neurological diagnosis and neurosurgery: CT, MRI, and DTI. Doi:10.103/npre.2009.3267.5

On the early CT units, the only technique of imaging available was what we now call axial mode or step-and-shoot. The later term better captures the rhythm of axial mode imaging since all the data necessary for a single slice is collected (shoot) in a spin before the patient is moved (step) to the next slice position. While axial mode has advantages in some circumstances and is still available on scanners, it takes more time than helical scanning since the stepwise movement of the patient is time-consuming relative to the time spent actually scanning. On early scanners with only a single detector row, the act of decreasing slice thickness by half would result in doubling the scan time. That is because scanning the same anatomy but with thinner sections was just like walking but taking smaller steps. The process of acquiring single axial scans had other limitations and many were due the relatively long scan time. For example, if there were any patient motion during acquisition of those single scans, misregistration or steps would appear between slices on reconstruction (Figure 1.10). This aversion to patient motion during axial CT scanning, imprinted on imagers for over a decade, made the spiral CT technique all the more remarkable when it was introduced in 1990. Now, patient

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0 0 0 0 0 0 0 01

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Mathematical 3 3 0 0 0 0 0 0 0 0 attenuation

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Figure 1.8 This illustrates just one pass of the collector and gamma ray source across a phantom containing water surrounding a aluminum rod. In CT language, this simple motion was called “translate.” After each pass across the object, the entire assembly would rotate 1 degree and collect another projection; so, this motion of first-generation CT scanners was called “translate-rotate.” The line below marked “mathematical” shows a numeric representation of the attenuation measurements collected by the detectors that could be used for image reconstruction. This information can also be represented graphically, as seen in the line “projection.” The first CT images were made using an algebraic reconstruction, but later all CT scanners used the projections in a reconstruction technique called back-projection, or more specifically filtered back-projection, because it proved to be faster than the purely algebraic reconstruction using computers of that era.

motion became a requirement for CT scanning. This innovative approach to CT imaging is credited to Willi Kalender, and the terms “spiral” and later “helical” were used to describe the path now traced by the rotating X-ray beam onto the moving patient (Figure 1.11). Helical imaging at fi rst was limited by scanner hardware, and only a short section of anatomy at a time could be covered in a scan because the wires that attached the X-ray tube to the gantry had to be unwound. Eventually, CT hardware was improved to maximize the benefits of helical scanning, and once continuous gantry rotation became possible, CT scan times dropped precipitously. Continuous gantry spin was made possible by the use of a slip-ring attachment between the tube and detectors to conduct power and data, respectively. But this was not a uniquely CT invention as slip rings were already commonplace on tank turrets and home TV antennas (Figures 1.12, 1.13). When we perform CT in the axial mode, the data for one slice goes on to image reconstruction as a discrete packet of information. In helical mode, since the X-ray beam actually sweeps obliquely to the moving patient, each of the axial slices must be created using data collected from more than one of those rotations. The attenuation values for the direct axial slice, or from any other plane for that

History and Physics of CT Imaging X-ray tube

Detectors

Figure 1.9 The arrangement of tube and detectors in a third-generation CT scanner. Unlike the “translate-rotate” approach, in this design, the tube and detectors move in a circle around the patient. While early versions of this design used a single row, current CT scanners use the same design but incorporate multiple detectors rows each with hundreds of individual detectors.

Figure 1.10 The irregular contour of this skull (arrows) is due to patient motion during the acquisition of the axial scans used for the reconstruction.

matter, are estimated from the known data points that were measured during the helical scan. This process of estimating the attenuation values in nearby tissue using the known, but only nearby, data, is called interpolation. It is really very much like the method used to estimate the value of a house before it is placed on the market. To provide a reliable estimate of a sale price, the appraiser does not actually add up the value of the many components of a house to determine its market value. The

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CT IMAGING (A)

(B)

Figure 1.11 In axial mode (A), the CT scanner gantry spins around just once and in the plane perpendicular to the patient. In helical mode (B), the gantry spins in the same but continuously while the patient table moves through the its center. The combination of these simultaneous motions (i.e., continuous rotation of the tube and the advancement of the patient), results in an oblique path of the X-ray beam across the patient. This X-ray beam trajectory can be described as “helical,” and this term is preferred instead of “spiral” since that term implies a continuously changing diameter as well.

projected selling price is based almost entirely on the recent sale prices of comparable houses nearby. For example, if there had been completed sales during the past year of the houses on either side of your house, the estimated value of your house would be much more reliable than if you were selling a custom built ten room mansion is upper Maine and the closest reference houses were in towns many miles away. The same principle holds true for helical imaging. The closer the helical wraps are together, the more accurate those estimated or interpolated attenuation values will be in tissues not directly in the scan trajectory. This explains why the use of a low pitch, that allows interpolation over a shorter distances, provides better resolution. In cases where a pitch value less than one is used, the overlapping of the helical sweeps allows the scanner to measure attenuation of some of the tissues more than once and that also decreases noise but at the expense of time and patient dose.

Multidetector CT: Beam Collimation Versus Detector Collimation The fi nal advance that will bring us up to date with modern CT scanners was the addition of multiple detector rows to the helical scanner. It is worth acknowledging that the very first EMI scanners also acquired more than one slice at a time, so the notion is not entirely new to CT, although the rationale for it changed with the different generations of scanners. On those very early translate-rotate scanners, a single rotation around the patient might take 5 minutes, so the use of a pair of detectors could significantly reduce total scan time. With the arrival of second and third generation designs, however, the second detector row was dropped presumably to save cost and reduce the complexity of reconstruction.

History and Physics of CT Imaging

Figure 1.12 Hard for many to believe now, but there was a time when the TV signal was collected free using a fixed antenna attached to the roof of a house. The quality of the TV image was of course related to the strength of the signal received and that meant, for many rural households far from the transmitters, that decent TV signal reception required sensitive antennas. The best of these could be rotated remotely from the living room while standing near the TV set in order to optimize the direction of the antenna and viewing the image as the antenna was rotated. By using slip ring contacts on the shaft of the antenna (arrows), the antenna could be rotated in either direction without worrying about later having to climb on the roof to unwrap the antenna wires. This was no small comfort on a cold Vermont winter night.

Twenty years after the EMI scanner, Elscint reintroduced the use of multiple detector row CT, but the rationale at that time was to limit tube heating during helical scanning. After the arrival of slip ring scanners, many sites were experiencing unwanted scanner shutdowns when performing wide coverage, helical imaging and that was because continuous scanning would make the X-ray tubes of that time overheat. Once that occurred, it required a forced break from imaging to provide time for the X-ray tube to cool off. This often occurred in inopportune moments, for example while imaging a patient after major trauma, and there were few precedents since it had been only rarely encountered previously when using CT scanners in the axial mode. This was because the time spent moving the patient between each rotation of the gantry, albeit short, provided enough time for the X-ray tube to cool off. Elscint’s design was intended to limit tube heating by decreasing the duration of the “tube on” time for the helical scans. Manufacturers quickly found there were other significant benefits of multidetector scanning, even after the tube heating problems were minimized by the introduction of X-rays tubes with substantially more heat capacity. While early multidetector scanners could provide either faster scan times or thinner slices, as the number of detector rows increased it became possible to provide both. Over the course of the next decade, scanners would appear with 4, 8, 16, 64, 128, and most recently 320 rows (Figure 1.14). Keep in mind that multidector arrays come at a cost since each detector row still contains nearly a thousand individual detector elements, and the use of metal dividers between rows to limit scatter meant that these multi-detector arrays are heavy, difficult to build, and expensive.

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Figure 1.13 A slip ring on a CT scanner (arrows). The contacts fixed on the large plate on the left ride on the circular conductive metal rails provide power and convey data while the entire gantry freely rotates.

Users need to be aware of exactly how the detector rows are arranged on their scanners since that can vary among the different manufacturers, and there is almost no way to know their arrangement intuitively. It is also important to recognize that some scanners provide fewer data channels than the number of available detector rows. So, a manufacturer may offer a scanner called the “Framostat 40” with only 20 data channels. In that case, you will fi nd that the scans can take longer than expected when using the thinnest detector collimation because only half of the total detector rows are active at the smallest detector collimation (Figure 1.15). The advantage of offering choices for the activation of detector rows is that it gives the user the options of using either the narrow center detector rows to provide the best detail or using all the rows for rapid coverage of large anatomic regions. So, keep in mind that your choice of “detector collimation” is not trivial since it determines not only the scan resolution but also the total number of detector rows activated, and that has a significant effect on scan time.

CT Image Contrast At the risk of stating the obvious, the shades of gray on a CT image are based on a linear scale of attenuation values. Wherever the X-rays are significantly absorbed or deflected, i.e. attenuated, by the tissues, very few X-rays will arrive at the detectors and those corresponding tissues will appear white on the image. Wherever there is little or no attenuation of the X-ray beam, more X-rays will arrive at the detectors and those tissues will be represented as black on the CT image. That is why air

History and Physics of CT Imaging

Figure 1.14 This fountain pen was placed on the plastic shield in this 320-detector row scanner to provide some perspective to its width. Using this scanner, the detector array is sufficiently wide to cover the entire head in a single rotation of the gantry.

appears black, bone appears white, and fat and brain are represented as shades of gray in between (Figure 1.16). This direct correlation of just the single value of X-ray attenuation with gray scale display differs substantially from magnetic resonance (MR) images, where there are multiple sources of information displayed on image, and so a dark area on the image might be attributed to signal loss from flow, low proton density, or even magnetic susceptibility effects depending on the scan technique and the anatomic location. Although CT imaging seems simpler than MR in principle, a number of factors confound our ability to assign the correct attenuation values to the imaged tissue and there are many illustrations of this problem included in the case fi les. For example, a renal cyst may appear to have higher attenuation on CT due to pseudo-enhancement (Chapter 8, pitfall 1), or CSF in the sella may be mistakenly assigned the same attenuation value as fat due to beam hardening (Chapter 5, artifact 6). So, while CT image display seems to be more straightforward than MR imaging, you must fully understand the many factors that can confound the accuracy of attenuation values displayed on a CT scan.

Slice Thickness While early scanners produced images with choppy images with visibly large pixels, since they used a matrix of 80 ×80, the in-plane resolution of CT images improved quickly. With each new generation of CT scanner, pixel size decreased fairly quickly to the current submillimeter standard size. But CT image resolution is determined by voxel size and that is determined by both the pixel size and the slice thickness.

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64 - 0.625 mm detector rows total widh = 4.0 cm

16 - 1.5 mm detectors

32 - 0.625 mm detectors

16 - 1.5 mm detectors

Total width = 7cm

Figure 1.15 These two different scanners both have 64 detector rows on their detector arrays but provide different usable scan widths depending on how they are activated. In the top example, the 64 detector rows are each 0.625mm wide, evenly spaced, and there is a data channel for each row. This arrangement could offer -.625mm detector collimation with a total usable scan width of 4cm. In the lower example, there are also 64 rows but only the center 32 rows are 0.625mm wide. The remaining 32 rows are all 1.5mm wide and arranged as a pair of 16 detector rows on the outside of the array. A scanner with this arrangement would offer only 2cm of coverage when using 0.625mm detector collimation, and that is half of that of the upper arrangement. However, using the center rows in pairs, they would function like an additional 16 1.5mm rows, and using that arrangement the total usable array width becomes 48–1.5mm detector rows. This would provide 7cm of coverage with each rotation, and that is nearly twice that of the upper array. So one manufacturer might offer their scanner with the lower arrangement to provide a wider array width for rapid body or lung imaging with the option to do finer imaging, like brain CT angiography. However, a CTA using a detector collimation of 0.625mm with the lower array would take twice as long as the same scan using the upper array. You need to know how the detector elements are arranged to correctly design scan protocols on your scanner for different imaging requirements.

Whenever images are created using thick slices, small structures may be obscured because each voxel is represented by a single attenuation value, and that is determined by the average attenuation of all the contents. This resembles the presidential primary process for states like Florida. There, all the delegates are awarded to the overall winner, unlike in New Hampshire, where they are fractionally awarded based on the candidate’s portion of the total vote. For example, if a single voxel contains both fat (low attenuation) and calcification (high attenuation), the mean attenuation of that voxel could be exactly the same as normal brain, making both the fat and calcification inapparent on a CT scan. It is more common to fi nd that a very small, dense calcification that occupies only a fraction of a voxel will cause the entire voxel to have the attenuation of calcium and that will result in an exaggeration of actual size of the calcification on a CT scan. On early single-slice CT scanners it was undesirable to decrease slice thickness for most imaging tasks because that significantly increased the time required to compete the scan. However, when

History and Physics of CT Imaging

Figure 1.16 This patient was lying on an ice bag during the CT exam performed for neck pain. Notice that the ice blocks (arrow ) are darker than the surrounding water. By CT convention, this means that the ice attenuates the X-ray beam less than liquid water. Since both the liquid water and solid ice have exactly the same molecular composition, this difference in attenuation must be the result of the slight separation of water molecules as water changes state to crystalline ice. In addition to this high sensitivity of CT imaging to differences in attenuation, it also provides sufficiently high resolution to show the air, note the dark spots, frozen within the ice.

using CT scanners with multiple detector rows, scan time is for all practical purposes independent of slice thickness. For example, a scanner with 64 channels using sub millimeter slice thickness can provide faster scans over comparable anatomy than can a four-slice scanner using 5mm slice thickness. This capability of multidetector scanners to provide very thin slice thickness without adding to scan time has made high quality multiplanar reconstructions commonplace.

Isotropic Voxels and Reconstructions The ability to scan with very thin sections has proved to be among the most significant advances of modern CT imaging. While early CT scanners were capable of providing good quality axial images when viewed slice by slice, whenever they were reconstructed into another plane of display the quality of these reconstructions was surprisingly poor due to thick slice thickness. For example, using a slice thickness of 1cm meant that the depth of each voxel was more than 10 times larger than the pixel size. These asymmetric voxels resulted in reconstructions with a striking “stair-step” appearance that were of little diagnostic value apart from gross alignment. However, the ability to scan using cubic or isotropic voxels in which the slice thickness is the same as the pixel size provides reconstructions in any plane that are equivalent in quality to the images in the acquisition plane (Figures 1.17, 1.18).

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Figure 1.17 These drawing show the difference between (A) voxels created using thick CT detector collimation, called anisotropic, compared with (B) those using very thin detector collimation, called isotropic voxels. Isotropic, or cubic voxels, are created when the slice data is nearly the same dimension as the length of one side of a pixel. For example, when using a 512 × 512 matrix for scan reconstruction, the detector collimation needs to be less than 1 mm in order to provide cubic voxels. The advantage of creating isotropic voxels is that the scan reconstructions in any plane (e.g., sagittal, coronal, or oblique) will be nearly equivalent in quality to images in the plane of acquisition. Illustrations provided by Dr. Rihan Khan, University of Arizona, Department of Radiology.

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Figure 1.18 Sagittal view made from standard 5mm reconstructions (A) and the 0.7mm original scan data (B).

While high-quality reconstructions are routine now for body and neuroimaging, it is important to consider that when using the thinnest available detector collimation, the signal-to-noise ratio (SNR) on each slice will be less than that available with the use of either wide detector collimation or slice reconstruction thickness when using narrow detector collimation (Figure 1.19). If the thin sections are to have the same SNR as thicker sections, the radiation dose for the scan must be increased. In practice, however, this problem is mitigated because the thin sections are rarely viewed primarily. By reconstructing the submillimeter data images in the desired plane of section at 3–5 mm slice thickness, the overall SNR is significantly better the thin source images. The principle of “scan thin, view thick” is the basis of most brain imaging because detector

History and Physics of CT Imaging (A)

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Figure 1.19 Notice that the noise visible in the 0.625mm section (A) becomes less apparent after merging data from multiple detectors together into a thicker slice, here as a 4.5 mm slice (B).

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Figure 1.20 Helical imaging requires collecting data from either 180 or 360 degrees of tube rotation so that corresponding views are available of any structure (note black structure A). However, off-center structures (note black structure B) may only be imaged once because of the divergence of the X-ray beam necessary for CT scanners with wide detector arrays. This undersampling artifact is called “partial volume” and it can results in blurring of the margins of that structure. This artifact should not be confused with volume averaging the (see Chapter 7).

collimation also minimizes beam hardening artifacts in posterior fossa and, for helical imaging, cone beam and partial volume artifacts ( Figure 1.20). But, when considering X-ray dose in this context, keep in mind that a small increase in dose can provide sufficient image quality for high quality reconstructions and that will ultimately save patient dose if it eliminates the need for a second scan. For example, by reconstructing axial CT data of the paranasal sinuses into the coronal plane, it eliminates the need for direct coronal scanning and thus reduces the total patient dose for the scan by nearly 50%.

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Image Reconstruction and Detector Arrays Hounsfield’s fi rst CT experiments used a pure algebraic reconstruction (Figure 1.21) to create images. In fact, it would appear that his first device was basically designed to collect the numbers in a manner best suited to solve the reconstruction formula. Although effective, algebraic reconstruction proved impractical for two reasons. First, it is very computationally demanding, and, second, it is impossible to use straight calculations to solve for the unknowns in an equation when the known values are not quite correct. That is the case with CT mathematical reconstructions since CT measurements include noise and a whole variety of artifacts. While there has been renewed interest in pure algebraic reconstruction techniques now that computers are fast enough to make it feasible, most CT scanners still use a less demanding approach called back-projection or more accurately filtered back-projection. The scan information can be thought of as a series of projections rather than a set of numbers (as shown previously in Figure 1.8). This technique, patented by Gabriel Frank in 1940, was originally proposed as an optical back-projection technique 30 years before the discovery of CT (Figure 1.22). To correct for the edge artifacts that are inherent with back-projection, an additional step is added to improve the quality of the fi nal CT images. This step is called filtering, although that term should not be confused with the physical act of fi ltering of the X-ray beam use to eliminate the low-energy X-rays. There are many filters, also called “kernels” which eliminates the confusion with metal X-ray filters, that the user can choose for reconstructing CT images. These range from “soft” fi lters that reduce noise at the expense of some image blurring to “sharp” fi lters used to display bone but increase apparent noise. The process of fi ltering occurs after data acquisition but prior to image display and can not be modified by the viewer afterwards. This of course differs from the setting of window and level used to view the reconstructed images (Figure 1.23). Upon the introduction of helical scanning, a new method for CT reconstruction was necessary to allow reconstruction of date acquired in a continuous fashion as the patient moved past the

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Figure 1.21 This 3 × 3 matrix demonstrates simply how one can use the sum of all the rows, columns, and diagonals outside the matrix to predict the value of the central, unknown, cell. In this simple example, the value of the blank cell in the center is of course 3. Early scanners used an 80 ×80 matrix that required hours of calculations using this algebraic approach. That approach was soon replaced with back-projection reconstruction techniques largely because they are faster.

History and Physics of CT Imaging

Figure 1.22 In this drawing from Gabriel Frank’s patent on back-projection, you can see that it was initially intended it to be a visual projection technique. Image B shows the light inside a cylinder that has collected the projections of the revolving object, line by line, in A. CT now uses a mathematical, not optical, application of this concept for reconstruction. From AG Filler. The history, development and impact of computed imaging in neurological diagnosis and neurosurgery: CT, MRI, and DTI. Doi:10.103/npre.2009.3267.5

detectors. This style of reconstruction incorporates the notion of estimation or interpolation of attenuation values for those tissues that fall between those actually measured during the X-ray beam sweep over the body (Figures 1.24, 1.25). Other challenges had to be addressed with each advance in CT technology complexity. For example, techniques needed to be developed for reconstruction of data collected simultaneously from each channel of a large multidetector array in helical mode. This was not simply an issue of handling larger datasets. As the number of detector rows increased, the X-ray beam increased in width in the craniocaudal direction to cover the array. That is why the thick fan beam of CT is sometimes called a “cone beam.” Since the beam arises from a small focal spot on the anode, the X-rays striking the outer detector rows arrive at a much steeper angle compared with those in the center rows. As a result, even for a uniform phantom, the X-rays arriving at the outer rows will have a longer path than those in the center. The already complex reconstruction algorithms now had to accommodate the differences in X-rays path lengths. As one might expect, these new methods for reconstruction also introduced some new and unfamiliar artifacts. The computational requirements for image reconstruction increased as the total number of detectors used for data acquisition exploded. Considering that since most scanners now have 700–1,000 separate detectors in each detector row, one rotation of the gantry provides a stunning amount of data to process. For example, one commercial dual-source scanner has over 77,000 separate detector elements in its two arrays that are intended to continuously collect data during each subsecond rotation of the gantry.

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Figure 1.23 These four images illustrate the difference between image filtering and windowing. The image in A is processed with a soft tissue filter and is displayed at a soft tissue window. The image in B shows the same dataset but now processed with a bone filter but displayed with the same soft tissue window and level as in Figure A. Notice how much more noise is apparent as result of this change in filter. The image in C was also processed using a bone filter but it is displayed with a bone window and level. Notice how much detail is now evident in the skull bones. The image in D shows the scan data displayed with the same bone window and level, but reconstructed using a soft tissue filter. Notice on this image how the bone edges appear much less sharp than image C. These paired images illustrate the balance between edge enhancement and noise that is determined by your choice of filter. Your choice of filter, also called kernel, will indirectly influence the dose necessary to scan the patient since it is an important factor in your perception of noise on the images.

Cone Beam Imaging A logical next step in the evolution of cross-sectional imaging would replace the complex multidetector array with a single flat detector similar to the ones that have replaced image intensifiers used for conventional angiography (Figure 1.26). While there are some similarities in configuration between a wide multidetector array and a flat-panel detector, there are also some significant differences to consider.

History and Physics of CT Imaging

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Figure 1.24 When using helical mode for scanning, the X-ray beam trajectory is angled to the long axis of the patient, and this angle increases as pitch increases. In order to assign attenuation values to the voxels that lie in between the actual beam path, the attenuation values need to be estimated or “interpolated” from known data points. And, the further away those directly measured points reside, the greater the degree of estimation. In this drawing the numbers that are not circled must be estimated based on known values determined from the directly measured points that lie on the oblique lines (solid lines).

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Figure 1.25 In this drawing there are more known values (circled) so there is less estimation of the values between the oblique lines necessary. Note the the values in between the solid circles are different from those in Figure 1.24. This illustrates why high pitch helical imaging, since the scan lines are farther apart, will have lower resolution.

Unlike conventional X-ray images, in which both direct and scattered X-rays contribute to the image, early CT scanners used a relatively narrow- ray beam that limited the contribution of scattered X-rays to the final image. As the number of detector rows in modern CT scanner’s detector array increased the beam became wide in two directions, its shape now resembled a cone rather than a fan (Figures 1.27, 1.28) since it must diverge from the anode in two directions, i.e., side-to-side and top-to-bottom. To minimize scattered X-rays from striking the detectors when using the wide fan beam in a usual multidetector scanner, the detector arrays incorporate thin metal plates, called septa, between each detector row. These septa absorb most of the scattered X-rays and are designed to allow only those X-rays oriented perpendicular to the detector to contribute to the image. While the use of septa improves image quality, they add weight and complexity to the array and also add to patient dose. A CT scanner using a flat-plate detector must also have a wide beam in two dimensions to provide even coverage of the flat panel. The terminology gets somewhat confusing, since the beam shaped used on a multidetector scanner can also be described as a cone beam, but many authors call any CT device using a flat-panel detector instead of multiple row detectors a “cone beam scanner.” But, these flat panel scanners, since the panel does not lend itself well to the use of septa common to multidetector scanners, must offer other methods to minimize the deleterious effect of scattered X-rays on image contrast. The use of a grid, not unlike those used with conventional X-ray films, can improve image quality but their use again requires an increase in patient dose. For example, as much as 20% of the total patient dose may be lost in the septa of a multidetector scanner and it is anticipated that this percentage could be more when using a grid on a flat panel or cone beam scanner.

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Figure 1.26 This image of an angiography unit during assembly demonstrates the flat-panel detector (arrows) at the top of the C arm with its X-ray tube at the bottom.

Figure 1.27 Multidetector scanners use an X-ray beam pattern that resembles the blades of this kitchen tool, used to cut butter into flour, since it also diverges in two directions.

History and Physics of CT Imaging X-ray source

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Figure 1.28 The usual third-generation CT scanner design has the X-ray tube (top, left image) move around the patient accompanied by the detectors (bottom, left image) that are rigidly attached opposite the tube on the gantry. Viewed from the side, the X-ray beam on a single-detector scanner is very narrow from head to foot and like a paper fan (middle drawing ). However, to accommodate the multiple detector arrays on modern scanners, the X-ray beam must be wide from head to foot as well as from side to side (far right drawing ).This figure provided by Josef Debbins PhD, Barrow Neurological Institute, Phoenix, Arizona

These factors are considered in the term dose efficiency, and this measurement is the composite of both the absorption efficiency and geometric efficiency of the scanner hardware. For example, the early single-slice scanners had a very high geometric efficiency since almost all the X-rays in the beam were collected by the single detector row. However, those early scanners had relatively low absorption efficiency because of the materials then available for the detectors. This has improved so that modern CT scanners offer a very high absorption efficiency, >90%, but a lower geometric efficiency compared with single-slice scanners. This give and take explains the surprising fact that the patient dose using a single-slice scanner in axial mode may be lower than the dose for an equivalent CT scan using modern multidetector scanner in helical mode. So, if dose increases and contrast decreases using a flat panel for CT, why bother? One reason is that flat panel scanners offer the potential for improved resolution compared with multidetector CT. Another is that a flat panel detector weighs considerably less than a large detector array and that offers the possibility of faster rotation times. But there is another limiting factor to rotation time that is rarely considered these days called recovery time. The limit to gantry rotation speed is usually considered to be the physical limits of spinning a very heavy object at high speeds. But another limiting factor is the time necessary for the detectors to reset after each exposure to the X-ray beam. For example, there would be no point in spinning the gantry at four rotations a second if it required a full second for the detectors to return to their baseline state after each exposure. This time necessary for the detectors to reset, also called afterglow, was a problem with older detector design but is negligible on modern multidetector scanners. However, flat-panel scanners will require more time for recovery so even though the gantry can physically spin faster, it won’t matter unless faster recovery time for the detector panel become possible. Dose constraints and potentially lower contrast, along with complex reconstruction algorithms have proved to be obstacles to the commercial development of cone beam CT for the time being. But this design does offer some advantages, and it deserves our continued attention since it is likely that

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many of these problems can be addressed with ongoing development of this technique. Cone beam CT is currently offered as an option on some angiography units and has proved to be useful in that setting for problem solving during complex interventions and the management of emergencies during interventional procedures.

Iterative Reconstruction All current CT scanners use variations of back-projection for image reconstruction. Recently, many scanner manufacturers began offering variations of mathematical or algebraic reconstruction, usually called iterative reconstruction (IR), for their scanners. There are two good reasons why. First, because of the increased utilization of CT, there has been an appropriate emphasis placed on reducing CT dose. Second, as a result of the relatively low price of supercomputer capabilities, it is now feasible to perform algebraic reconstructions at acceptable speeds and cost. Early indications suggest that dose reductions on the order of 50–75% are feasible for body imaging using IR without significant compromise in image quality using these mathematical reconstruction techniques. Many variations on this theme are now provided by vendors of CT equipment. Some versions even limit noise by accounting for the specific errors in the imaging chain, also called optics. Others, rather than use a pure mathematical reconstruction, use hybrid techniques that start with the traditional fi ltered back-projection but then use a mathematical technique to reduce noise by comparing that reconstruction to the raw data in an iterative process. The term “iterative reconstruction” describes a process of revising the image data in order to provide a “best fit” with the actual scan data. This is done in a continually updating, or iterative, process. I think of this much like the way one fills in a crossword puzzle (Figure 1.29). The reason most of us use a pencil to fill out these puzzles is because we may find opportunities to reconsider our response to an “across” clue once we figure out the “down” clue in that same location. I think of the iterative reconstruction process in this simple way; the software takes it best shot at creating the image, then goes back to the raw data to see how well it did, adjusts a few things, and checks again to see if that fits any better. One reason why this cannot be easily accomplished in a single, powerful calculation is that the raw data itself contains errors and noise. As a result, there is no single solution for the calculations, and so the most that can be hoped for is the creation of a “best fit” for that dataset. Think of it like a crossword puzzle but, in several spots on the grid, there is no word can satisfy both the “across” and “down” clues. While IR can be used to reduce dose or improve image quality at the same dose, it does require special software and computer hardware and currently it adds time for processing. Nevertheless, because it holds considerable promise for significant dose reduction and will be widely adopted in some fashion. This approach also offers new tools for minimizing streaks that arise from implanted metal. While some other and less expensive postprocessing options are available that do not refer back to the raw data in the same way, these should be considered carefully since they present the risk of creating “pretty” images at the expense of smoothing over clinically important contrast. For example, a postprocessing algorithm that eliminates noise in homogeneous areas of anatomy could potentially obscure true but subtle differences in attenuation. But iterative reconstruction combined with large decreases in dose will without doubt have its own limitations, and it will take some time to validate all these new techniques in the clinical arena before they can be used with complete confidence.

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Gantry Angulation and Image Display Most single-slice CT scanners included a mechanism to tilt the scanner gantry relative to the patient table. This was used on a regular basis to optimize the plane of imaging for axial brain scanning or, when combined with head tilt, to provide direct coronal images of the brain or sinuses. One substantial benefit to angulation on early scanners was that, by using tilt, one could minimize the number of

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slices necessary to cover the brain. And when imaging with CT was considered in terms of “minutes per slice,” eliminating one slice was not trivial. As scanner speed improved, the primary function for angulation in brain imaging was dose reduction to the eyes, and it was generally recommended to exclude them from the scan since they are susceptible to radiation injury. Now, however, on most scanners in helical mode and those units with large multidetector arrays or two sources, gantry tilt is not available for brain imaging. In spite of this change in hardware, it is commonplace to continue to present head CT scans with the traditional angulation since it is familiar to imagers and it makes comparison with prior CT scans easier. While gantry tilt had been used with patient positioning to provide direct coronal imaging for temporal bone and paranasal sinus imaging, since most modern scanners offer near isotropic voxel images, direct coronal imaging is really no longer necessary. Now, even reconstructions in sagittal views that were formerly unthinkable are routine. In fact, isotropic voxel imaging has created an imaging environment that resembles MR since even oblique reconstructions of diagnostic quality are now available on multidetector scanners in both axial and helical modes (Figure 1.30A and B). The loss of gantry angulation has created two new problems, however. The radiation dose to the eye is lowest on those scanners that offer gantry angulation if the user prescribes the scan angle and range to exclude the orbits. However, on scanners that do not allow gantry angulation, the eyes are always included in the scan but the imager may not be as aware that the eyes were included if the data is reconstructed into the traditional display angle. So, while the lens is always included on head scans performed on new scanners without gantry angulation, the measured dose to the eye during direct helical imaging with a modern multislice scanner is still quite low. This represents another one of the compromises of CT imaging. As scanners enlarged to incorporate multiple detector rows, the tilt option was lost but the potential for increased dose was offset by more sophisticated automatic exposure control, beam filtering, and diminished dose from overbeaming with more detector rows (see Chapter 2, Overbeaming). While the use of automatic exposure control for brain imaging may not make sense otherwise for a roughly (A)

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Figure 1.30 A, B This high quality coronal CT image (A) was reconstructed from the thin section axial imaging data. Note the small defect in the bone of the sphenoid sinus (arrow) that corresponds to the site of a CSF leak noted on the coronal MR T2 weighted scan (B, arrow).

History and Physics of CT Imaging

spherical object, it can be worthwhile by providing greater dose reduction to the lens. Another option to reduce lens dose is to use bismuth X-ray attenuating eyecups, but this adds cost and time (see Chapter 2, Shielding). The second problem encountered with brain scans performed without gantry tilt is that the user needs to be attentive to artifacts from hardware in the mouth, such as amalgam, crowns, and implanted posts. While these were almost never an issue when gantry tilt was used, the metal artifacts arising from X-ray shadowing behind these very dense materials frequently projects directly over the posterior fossa and, in some cases, significantly degrades the diagnostic value of the CT scan (Figure 1.31). One option to minimize this artifact is to instruct cooperative patients to tuck their chins during the scan. This recreates the traditional imaging angle without requiring gantry angulation and should be helpful in limiting the metal artifacts from teeth and, if carefully done, it offers the potential for reducing eye dose as well. Medical practice is at times an odd mix of eager acceptance of new technology and rigid resistance to change in almost every other way. With the arrival of scanners without the capability of gantry angle, the only real benefit now to viewing CT brain scans in the old fashion is that the orientation is familiar to imagers. Straight imaging in many respects would make it easier to compare CT scans with MR scans, since the later are routinely displayed without angle (Chapter 5, Artifact 9). But it seems likely that, as more centers move to isotropic imaging of the brain, head scans will eventually be presented in two or three orthogonal planes for review similar to the way most body CT images are displayed now.

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Figure 1.31 The axial CT scan (A) shows considerable artifact overlying the cranio-cervical junction without a clear source. The coronal reconstruction (B) shows that the streaks are arising from dental amalgam and projects over the skull base in this case because no gantry tilt was available on this scanner.

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Scan Acquisition Speed One of my friends with a very large front lawn once told me that he bought a new mower about every 10 years, and every new mower had a wider blade than the last. He figured that by the time he retired, he would be able to mow his whole lawn in one trip down and back. That is essentially how CT detector arrays have changed over time. While the earliest scanners had one or two detectors and took up to half an hour to scan the brain, there are now scanners with 320 detector rows that can cover the entire brain in a single, subsecond rotation. The three factors that determine scan speed using helical technique are beam collimation, table speed, and tube rotation time. The term “pitch” is a useful concept that incorporates all three factors into one term. Pitch may be defi ned in several ways, but we will use it specifi cally to mean: Table Movement in cm During Each Tube Rotation ÷ Beam Collimation in cm (2). For example, if the beam collimation is 4cm, the table moves 4cm per second, and tube rotation time is 1 second, then the pitch for that scan would be equal to 1. At an intuitive level, a pitch of 1 can be visualized as though the X-ray beam paints a continuous helical stripe around the patient, with no gaps and no overlap. Now, if we just decrease the rotation time to 0.5 seconds but keep the beam collimation and table movement the same, the pitch is now equal to 2. With that pitch, there will be wide gaps between the stripes covered by the X-ray beam, and that will decrease resolution and increase minimum slice thickness. To its credit, however, higher pitch values can be used to decrease radiation dose and overall scan time (Figure 1.32).

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Figure 1.32 This illustration shows how the use of a high pitch value (A) leaves wide gaps in the beam trajectory compared with a scan using a lower pitch (B). The advantage of using higher pitch values is that it offers faster scan times and potentially at a lower dose at the expense of resolution and minimum slice thickness.

History and Physics of CT Imaging

Rotation Time When scanning in helical mode, decreasing rotation time will have a significant impact on total scan time. Decreasing rotation time by half effectively cuts scan time in half and this is important for imaging moving tissues like the heart (see Chapter 3). When using step-and-shoot, however, the time spent actually scanning is just a fraction of the total scan time since that is disproportionately composed of moving the patient from slice to slice. Decreasing rotation time usually reduces overall image quality because fewer projections are obtained, but most new scanners routinely use rotation times of 0.5 seconds for CT angiography (CTA) since is the benefits of fast scanning outweigh the incremental image degradation.

Dual-Energy, Dual-Source CT Scanning Computed tomography scan times are remarkably short now, with a brain scan requiring less than 10 seconds. That can be compared with the 5 minutes per slice that was routine at the start of my radiology career. By my calculations, that means that a brain CT scan is about 250 times faster now compared with the same scan 30 years ago. I have wondered why the charge for CT has not gone down, too, but that is another matter entirely. One would have thought that CT scanners are now fast enough for any diagnostic problem, but faster scanning has created many new opportunities for CT imaging but a few problems, too. For example, when imaging the brain and neck, it has become necessary to build in a time delay after giving contrast because otherwise the scan will be over before sufficient contrast can accumulate in abnormal tissue (see Pitfall 8, Chapter 6), and CTA is now so fast that incomplete fi lling of the carotids can create an artifact easily mistaken for dissection. But these limitations should be recognized and adjustments made. Overall, faster imaging has created entirely new applications, and it seems that CT can never be fast enough for some purposes. One current approach to further decreasing scan time is to build scanners with two tubes and two detector arrays set 90 degrees apart (Figure 1.33). Because the attenuation value of an X-ray passing through the body should be about the same from left to right as right to left, only 180 degrees of imaging, or one-half of a rotation, is really necessary. Therefore, by using two tubes and detectors simultaneously, only a 90-degree gantry rotation is really necessary to acquire the necessary data for image reconstruction. Another advantage of using two completely separate and independent tube–detector sets on the gantry is that they can be set up differently. For example, if one tube operates at a low kV and the other at high kV, the same slice can be reconstructed from a pair of images with different image contrast characteristics. The difference in the way the tissues attenuate the X-rays at the two energies can then be used to better characterize the tissues. This technique, called dual-energy CT scanning, is particularly helpful for differentiating bone from contrast since they can have the same attenuation value when imaged using a single kV. That is because iodine selectively absorbs low-energy X-rays, called the photoelectric effect, while bone largely scatters X-rays at both energies. Since iodine will appear to have a much higher attenuation value when imaged using a low kV, these two,

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Figure 1.33 The arrangement of X-ray tubes and detectors in a dual-source CT scanner. While it still uses standard third-generation geometry, it requires only a 90-degree rotation to collect the equivalent of 180 degrees of rotation on a single-source scanner.

otherwise indistinguishable, tissues can be separated on the basis of their behavior at two different kV values. This has been shown helpful for many things, including differentiating a hemorrhagic renal cyst from an enhancing tumor, blood from contrast in the brain, and for creating CTA and CT venography (CTV) images with the bone removed (Figure 1.34). Dual-energy scanning is not just limited to scanners with two tube–detector sets. It can also be performed on some scanners that have a specially designed, but single, X-ray source that can rapidly switch between two different tube voltages during the scan. Because of the rapid image acquisition available on current CT scanners, several new imaging techniques have become routine, such as cardiac imaging and whole-brain perfusion, and there is some hope for time-resolved CTA in the near future. That would be of value in some cases because it is commonplace to see both veins and arteries on normal CTA. As a result, they cannot be used for the diagnosis or follow-up of small arteriovenous malformations of the brain or spine in most cases. If CTA scans can be acquired in distinct vascular phases (i.e., arterial, capillary, and venous), that information may be sufficient to fi nd areas of arteriovenous shunting noninvasively. The current limitations include insufficient temporal resolution on most scanners and the expected radiation dose incurred by repeated scanning over time, but these may be addressed in the near future.

CT Imaging Techniques As you can appreciate from this review, the hardware and software used for CT imaging are complex, both separately and in their integration. Variations in scanner design among different manufacturers make it very difficult to export techniques used on one scanner to another. For example, if the detector array on one scanner is farther away from the tube than another, it may require more

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Figures 1.34 A–E These axial images (A and B) were obtained simultaneously on a dual-source, dual-energy CT scanner immediately after intra-arterial thrombolysis of a middle cerebral artery occlusion. The kV 80 image (A) shows high attenuation material in the right basal ganglia and its measurement within a cursor shows an attenuation value of 349 HU. The corresponding image using a kV 140 (B),at the same location, shows an attenuation value of only 189 HU. The doubling of attenuation value between a low kV and high kV image is due to photoelectric effect and is characteristic of iodine. CT follow-up confirmed that the high attenuation in this patient was due to contrast staining. This carotid surface reconstruction seen in C illustrates one benefit of bone removal CTA when imaging at the skull base. By removing the bone surrounding the internal carotid artery, using its imaging characteristics at the two energies, this carotid stenosis is quite evident. It might easily be overlooked on conventional CTA since the surrounding bone obscures it on both the source image at that level (D, arrow ) and axial maximum intensity projection (MIP) (E) reconstruction.

tube current to provide equivalent images. That is, of course, unless it has more efficient detectors or thinner septa in the detector array. And we still haven’t accounted for the filtration devices used to modify the X-ray beam, and these can differ in both materials and shape. In fact, each manufacturer makes choices for the detector array, software, X-ray tube, and the like that, in sum, define

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the performance characteristics of their unit. So don’t be frustrated when the technique that always works on your GE scanner will not provide comparable images or dose on your Siemens unit. That is also why specific techniques will not be recommended in this book. What makes it even more complex these days to design CT protocols is that the manufacturers often build into their scanners automatic adjustments that may unobtrusively offset changes made by the user. The manufacturers have taken some control of the unit away from the user with the intent of ensuring good quality-images and patient safety. For example, you may fi nd that when you increase pitch to lower the patient dose the scanner has automatically increased the mA to offset the increased image noise expected from using the higher pitch value. This is most important to keep in mind when using automatic exposure control or AEC. In that mode, the scanner will adjust the dose to meet preset values of image “noise index” or “reference mA” equivalent. If this is the case, you will fi nd that “automatic” does not necessarily mean “lower.” Patient dose can reach surprisingly high values in large patients or when AEC is used inappropriately (e.g., for brain perfusion). And keep in mind that the reference values for AEC that are preset by the manufacturer may be ideal for image quality, but completely acceptable imaging can usually be achieved by lowering those reference values and thereby limit dose for your patients. You must be aware of how your scanner will respond to changes in technique and patient size if you are to truly manage X-ray dose. The antilock brakes or dynamic steering controls available on most new cars are intended to keep you out of trouble in challenging driving situations. But, by interpreting your steering and braking input, they separate you from direct control of the vehicle. On Ferraris (or so I am told), a switch on the steering wheel allows the driver to completely turn off these driving aids in order to deliver complete control of the car to the hopefully skilled driver, thus allowing them to explore the unfettered performance limits of the vehicle. Most CT scanners have no such switch, so the user will need to be aware that the software installed by the manufacturer may modify user input. Make it a habit to check that the scanner has not offset some change that you made in the protocol with an adjustment somewhere else. And that brings us to Chapter 2, on CT dose and dose reduction techniques. SUGGESTED READING Kalendar WA, Wolfgang S, Klotz E, Vock P. Spiral volumetric CT with single-breath-hold technique, continuous transport, and continuous scanner rotation. Radiology. 1990;176(1):181–183. Gupta R, Cheung AC, Bartling SH, Lisauskas J, Grasruck M, et al. Flat-panel volume CT: Fundamental principles, technology, and applications. Radiographics. 2008;28:2009–2022. Tan JSP, Tan KL, Lee JCL, Wan CM, Leong JL, Chan LL. Comparison of eye lens dose on neuroimaging protocols between 16- and 64-section multidetector CT: Achieving the lowest possible dose. AJNR.2009;30:373–377. Mahesh M. The AAPM/RSNA physics tutorial for residents. Search for isotropic resolution in CT from conventional through multiple-row detector. RadioGraphics. 2002;22:949–962. Karcaaltincaba M, Aktas A. Dual-energy CT revisited with multidetector CT: Review of principles and clinical applications. Diagn Interv Radiol. 2011;17:181–194. Goldman LW. Principles of CT: Radiation dose and image quality. J Nucl Med Tech. 2007;35:4:213–225. Bauhs JA, Vrieze TF, Primak AN, Bruesewitz MR, McCollough CH. CT dosimetry: Comparison of measurement techniques and devices. RadioGraphics. 2008;28:245–253.

History and Physics of CT Imaging Parry RA, Glaze SA, Archer BR. The AAPM/RSNA physics tutorial for residents. RadioGraphics. 1999;19:1289–1302. Barrett JF, Keat N. Artifacts in CT: Recognition and avoidance. RadioGraphics. 2004;24:1679–1691. Kilic K, Erbas G, Guryildirim M, Arac M, Llgit E, Coskun B. Lowering the dose in head CT using adaptive statistical iterative reconstruction. AJNR. 2011:32:1578–1582.

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RADIATION SAFETY AND RISKS Alexander C. Mamourian and Josef P. Debbins

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Patient Dose and Dose Reduction Techniques For anyone involved with medical imaging, it is essential to be on familiar terms with radiation terminology. Those who are not will fi nd themselves on unsteady ground either when reading the literature or discussing the particulars of a computed tomography (CT) scan with patients. The numerous terms used to describe radiation dose contribute substantially to the confusion, and I believe the fi rst step to a solid understanding is to use only two or three terms for dose measurement. Otherwise, you will eventually get lost by using the full range of units e.g. Roentgen, rad, rem, Gray, Sievert, and their respective conversion factors. In this book, the terminology of radiation dose will be limited to the most commonly used measurements. You will also fi nd it helpful to keep a few hard numbers in mind, such as background radiation effective dose (3mSv) and the American College of Radiology (ACR) limits in units of CT dose index volume (CTDIvol) for an abdominal CT scan (25mGy) and head CT scan (75mGy). These can provide reference points as you consider variations of technique. You can then build on that foundation over time as new techniques for dose reduction are introduced and your depth of understanding increases.

Commonly Used Measures of CT Dose CTDI VOL A number of CT dose measurements have been used over time, and these are variations of a term called CT dose index or CTDI. Some that are no longer in common use are CTDI FDA, CTDI100, and CTDIw. Fortunately, there is only one you should be familiar with now, the measurement CTDIvolume (CTDIvol)since it is calculated on most commercial CT scanners and is widely used in the literature. The meaning of CTDI has changed over time in response to advances in CT technology. The measurement of CTDIvol is performed in a phantom and incorporates some weighting of the peripheral dose compared with the lower central dose, as well as the pitch used for helical scanning. Keep in mind, however, that CTDIvol for a CT scan is just a calculated value based on X ray tube settings. Although it does account for the specific features of the scanner, such as beam filtration, when reported by the scanner at the end of a scan, it is not at all a direct measurement of the dose delivered to a particular patient. It is in truth a measurement of the X-ray tube output and is based on specific scan parameters, including mA, kV, and pitch, as well as dose modulation. It can be quite helpful as long as one recognizes that in some situations it may not accurately reflect the actual patient dose, and CT brain perfusion has been cited as one example. Although CTDIvol is by no means a perfect measure of dose it is useful in day-to-day CT operations. Many fi nd it helpful whenever they want to see how the modification of one scan parameter will influence patient dose because that change will be reflected in the CTDIvol. This number also allows comparison of the dose of the same scan performed on different machines even when they are from different manufacturers. And, by just glancing at the dose report, now available for each

Radiation Safety and Risks

patient exam, one can determine whether a specific exam falls within both the ACR’s and your own institutional guidelines. For example, if the dose page from a head CT scan shows a CTDIvol of 90mGy, it should serve as a warning that there is a problem with the scan protocol that should be addressed immediately.

Absorbed Dose The absorbed dose is the measure of ionizing radiation deposited into a specific volume of tissue. Although absorbed dose does not take into account the behavior of that tissue or provide an estimate of the risk of radiation, it does correlate well with the immediate or deterministic effect s of radiation. The units of absorbed dose in common use are the Gray (Gy) and milli-Gray (mGy) named for the physicist Louis Harold Gray. The actual absorbed dose can be quite difficult to determine since it depends on the energy of the X-rays, their total number, and the nature and depth of tissues involved. Although it is easiest to measure dose at the skin surface, it is the total absorbed dose that is of interest, and so this number for CT is always an estimate. The absorbed dose can be used to predict whether the radiation was likely to be the cause of reddening of the skin or even hair loss in a particular patient. Just so you have some benchmark values, keep in mind that the absorbed dose of a head CT is about 50mGy and temporary hair loss occurs at a dose of 3Gy at the skin. Thus, it would take over 50 diagnostic head CT scans to cause temporary hair loss, assuming there is a cumulative effect. There are reports of hair loss after a single CT brain perfusion scan in patients who also had more than one cerebral angiogram, thus supporting the concept of cumulative effect of diagnostic radiation.

Deterministic Effects Radiation sickness leading to death is the most severe of the immediate or deterministic effects of radiation and is usually associated with nuclear warfare or reactor accidents. There are rare exceptions, however. For example, about 5 years ago, Alexander Litvinenko, a former officer in the KGB, died from radiation sickness after being intentionally poisoned with polonium 210. His cause of death was initially overlooked since polonium kills with alpha particles, not gamma rays, making it very difficult to detect if this poison is not suspected. In clinical practice, deterministic effects are only rarely encountered because the absorbed dose of a diagnostic scan is so far below the threshold for effects. However, it is not uncommon to see skin and hair changes from radiation after prolonged interventional procedures or in patients who have multiple high-dose imaging procedures during a single hospitalization (Figure 2.1). In addition, it is very important for anyone involved with CT imaging to be aware of the possibility of deterministic effects and know what they look like. For example, it was reported that some patients who experienced hair loss after improperly administered CT scans were initially sent to dermatologists because the connection between the recent CT scan and their symptom of hair loss was not initially recognized.

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Figure 2.1 Photograph of the typical band-like hair loss from brain CT perfusion. In this case, it was temporary and secondary to the cumulative dose from multiple high-dose imaging procedures and not from the radiation dose incurred during a single perfusion scan.

Effective Dose The effective dose is measured in units of Sieverts or milli-Sieverts (mSv), named for physicist Rolf Sievert. It is a measurement of dose that, unlike absorbed dose, takes into account the sensitivity of the tissues that receive the radiation. For example, while CT imaging of the head requires a relatively large radiation dose, its effective dose is only 2mSv. That is less that the average individual’s effective dose from background radiation in the United States (3mSv). The reason the effective dose of a head CT is so low is that the brain is relatively insensitive to radiation. Compare that with a chest CT that has a much lower absorbed dose but an effective dose that is three to five times more than a head CT. That is because the organs included in the chest CT, such as breast and esophagus, are much more sensitive to radiation, and this is reflected its high effective dose.

Stochastic Effects The effective dose is use to predict the late or stochastic effects of radiation. The word “stochastic” is not commonly used in medicine apart from this circumstance. Its meaning in this situation is usually taken to be “randomly determined” but its derivation is attributed to the Greek word skokhos, “to aim,” although some would say more aptly from the word stokhazesthai, “to guess.” The magnitude of the stochastic effect for diagnostic radiation remains unclear, but it is generally accepted that the late adverse effects of diagnostic radiation may appear years after radiation exposure. What remains uncertain to this day is the precise level of radiation exposure necessary to even entertain

Radiation Safety and Risks

the possibility of these late effects. The authors of an article on virtual colonoscopy using CT wrote, “ An estimated 70 million CT scans are performed in the U.S. every year, up from three million in early 1980s, and as many as 14,000 people may die every year of radiation induced cancers as a result.” A more recent paper in the Lancet suggested that just three CT scans of the head incurred prior to age 15 will increase the risk of brain cancer in that patient by several fold. These are just estimates and difficult to confi rm since it is almost impossible to determine which cancers are related to radiation, among the many that occur every year. There is little doubt, however, that late effects can occur from a long exposure to high doses of radiation. It is worthwhile for anyone involved in medical imaging to read about those who died from radiation as a result of early medical or scientific investigations. Perhaps the most famous example is Marie Curie. She is credited in some share with introducing the term “radioactivity” and with the discovery of polonium (named after her homeland Poland) and, later, radium. Her tireless work to isolate the element radium required long exposures to pitchblende, a radioactive ore. She died in 1934 from aplastic anemia that seems undoubtedly the result of her prolonged exposure to radiation. The American Association of Physicists in Medicine (AAPM) in December 2011 took the position that, for diagnostic tests with an effective dose of 50mSv or less, when incurred in a single exposure, or 100mSv in multiple exams over a short time, the late effects may be “nonexistent” (their word). It is also important to keep in mind that medical radiation has a purpose, and the expectation is that the risk should be offset by the benefits of the exam for the patient. Since the actual risk of low-dose radiation is unknown, and a small risk is reasonable for a large gain, it is very difficult to know how much concern should be attached to a single CT exam. However, common sense would indicate that less radiation is better, and many patients are having multiple CT scans. These are the best reasons why medical imagers must be attentive to using the smallest amount of radiation necessary to make the diagnosis using CT or any other X-ray device. This idea is captured in the term ALARA (As Low As Reasonably Achievable) and should be our goal whenever considering the necessary dose for medical imaging.

Dose Length Product Conversion to Effective Dose Dose length product (DLP) is expressed in the units of mGy-cm. Its determination is really quite uncomplicated; it is simply the calculated CTDIvol multiplied by the length of the scan, from head to foot, in centimeters. But this simple calculation has been demonstrated to provide a remarkably accurate measure of effective dose when compared with the much more sophisticated Monte Carlo simulation. Although sounding very James Bond-ish, this method uses a mathematical representation of the body with approximate shapes of organs and their estimated radiation sensitivity to predict the effects of radiation. It has been demonstrated that multiplying the DLP by a predetermined constant for a specific body part will provide a value for effective dose in mSv that is a surprisingly close to the calculated effective dose using the Monte Carlo simulation. For brain imaging, that constant is .002 mSv/mGy-cm. Since the average head CT has a CTDIvol of under 70mGy and a DLP of about 1,000 mGy-cm, the calculated effective dose of a head CT is 2mSv (.002mSv/mGy-cm × 1,000mGy-cm). Now, if we were to use the exact same imaging parameters for

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a CT scan of the neck and assume it has the same DLP of 1,000 mGy-cm, we would then use a larger constant (.005 mSv/mGy-cm) to predict effective dose because of the increased sensitivity to radiation of the tissues in the neck. The conversion factor for chest CT is even higher, .018 mSv/mGy-cm, because the tissues there are even more sensitive to the effects of radiation.

Techniques for Dose Reduction ESTIMATING OPTIMAL DOSE One of the most difficult challenges for the diagnostic imager is to determine the “optimal dose” for CT imaging. This is in part due to the very subjective nature of this assessment, which can vary widely among individuals based on their experience and expertise. It is important to keep in mind at all times that the goal is not just to produce exquisite CT images. It should be to use just enough radiation to provide images as good as they need to be in order to establish the diagnosis. And, keep in mind that when we talk about “radiation cost,” it is the patient paying the price. That is why the CT user needs to find a reasonable compromise between patient dose and image quality. Although it is tempting to create images of extraordinary quality—and even if you find them easier to interpret—if this requires a higher dose, it must be justified by the clinical question. Another problem to consider when determining optimal dose for CT is our inability to perceive when a CT was performed using too much dose. Early X-ray images were no different from photographs taken with a fi lm camera. Those early fi lm negatives accurately refl ected the relationship between the available light, the sensitivity of the fi lm, and the camera settings. For example, if the aperture was set too wide on a sunny day, the picture would appear dark from overexposure. There was a time when making a high-quality X-ray image, in much the same way, required a good deal of judgment so that the proper settings of kV, mA, and exposure time were combined to give the correct exposure of the radiographic fi lm. I still remember my discomfort when looking at the nearly black X-ray image that was taken of one of my children by a student radiographer many years ago. Both of us clearly recognized that too much radiation was used. As many newly unemployed newspaper photographers well know, photography using digital cameras is much more forgiving because only a loose relationship remains between the exposure settings and the fi nal image. That is due largely to the use of electronic detectors instead of fi lm in modern cameras. These can provide acceptable images under a wide range of lighting situations just by modifying the sensitivity of the sensor. Within reasonable limits of X-ray dose, modern CT scanners function the same way as digital cameras since they also provide acceptable images within a wide range of exposures. In fact, when CT scans are performed with too much radiation there will be few complaints from imagers because the CT images actually look better—the higher the dose, the better the images. Modern CT scanners therefore require that the user have a good understanding of the dose reports since image assessment alone can be misleading. Even when CT images are made with too little dose they might pass for a time as adequate if their poor quality is attributed to other confounding factors like motion (Figures 2.2, 2.3).

Radiation Safety and Risks

Figures 2.2 This scan was mistakenly performed using a CTDIvol of only 19mGy, a value that is one-third of the usual dose used on this scanner. Although there is some degradation in image quality, it may be difficult to appreciate until you compare it with a scan of a different patient performed with the usual dose at CTDIvol of 60mGy (see Figure 2.3).

Figure 2.3 Scan performed with usual CTDIvol of 60 mGy.

CT Dose Reduction: Indications Years ago, a writer reflected that the most effective way to minimize travel time was to just stay home. In the same way, the simplest and most powerful way to minimize patient radiation dose is to not perform a CT scan at all. You must consider whether the patient will benefit from imaging and, if so, if CT is ideal for the diagnosis (since there are now many other options for imaging, such as magnetic resonance [MR] and ultrasound). Another strategy that immediately reduces dose, and one you may have more control over, is to severely restrict your use of high dose techniques such as multiphase CT. In most circumstances, there is no need for both pre- and postcontrast imaging of the brain or chest, for example. For sites that have access to a dual-source, dual energy CT scanner a technique called “virtual non-contrast” may be an option for dose reduction. For example, when obtaining a contrast enhanced brain CTA there can be substantial dose reduction for the patient by not performing a non-contrast head scan prior to the CTA. For example, with conventional CTA is difficult to determine whether there exists subarachnoid hemorrhage without a non-contrast exam. But, by using the imaging characteristics of iodine and blood at the different energies it becomes possible to provide a “virtual” non-contrast scan from the CTA data by removing the contrast enhancing structures (see Chapter 6, pitfall 5).

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Critical evaluation of the need for CT scanning does not require canceling all exams, just a thoughtful approach and some humility when considering whether our patient’s health has improved in proportion to the increased use of CT over the past 15 years.

Axial Versus Helical Scans An important starting point when designing a scan protocol is the decision of whether to use axial or helical mode imaging. In general, axial imaging will prove to be slower and have inferior reconstructions compared with helical mode, but, at the same time, it offers better control of the dose distribution and, in many cases, produces fewer artifacts. Other factors may not be immediately apparent as you consider these two options for brain imaging. For example, the use of helical mode means that the gantry cannot be tilted on some scanners. But for those CT scans that must be fast, such as chest imaging and CT angiography (CTA), helical mode is the better choice. With regard to dose distribution, there are two terms you need to understand when considering the choice between axial and helical: overbeaming and overranging.

Overbeaming and Overranging On all CT scanners, the shape of the X-ray beam that emerges from the tube is modified using metal plates called collimators. On the old single-slice CT scanners, the fan beam width was the same or less than the width of the detector row (Figure 2.4) and so the entire X-ray dose was utilized in some fashion to create the image. If we choose to discount differences in detector sensitivity, this arrangement provided a very high “dose efficiency.” This is not the case in multidetector scanners, where the beam width must always be wider than the detectors. That is because the X-ray beam must cover all the active detector rows evenly. Since the X-ray beam diverges from a point on the anode, a portion of the X-ray beam must always fall outside the end detectors in the array, and this dose is effectively wasted since it does not contribute to the image. The portion of the beam that extends beyond the detectors is called the penumbra, and this effect is called overbeaming. It is important to recognize that this extra dose occurs over the entire length of the detector array (Figure 2.5) and, on average, this adds about 1.5mm of extra tissue radiation on either side of the array. While seemingly small, this extra radiation adds up quickly when using a scanner with a narrow detector array since it will require many rotations to cover the chest and abdomen, for example. However, overbeaming can be discounted for scanners using 32 rows or more since they require so many fewer rotations to cover the same anatomy. As the number of detector rows increased, however, another source of added dose became more significant called overranging (Figure 2.6). Because of the very nature of helical image reconstruction, the X-ray beam must begin and end its path outside the region of interest. This bit of extra scanning is necessary to provide the data points needed for interpolation on the end slices. This added dose, unlike overbeaming, is not really wasted since it is necessary for reconstruction, but it is easy to overlook when considering patient dose. The magnitude of this added dose from overranging increases with beam collimation, number of detector rows, and the pitch value.

Radiation Safety and Risks X ray tube

Single detector row

Figure 2.4 On a single-slice scanner, the beam collimation is always the same as or less than the width of the detector row. Since all the X-rays are directed at detector elements, this arrangement is described as having high dose efficiency.

X ray tube

Figure 2.5 This illustration shows the added dose that falls on either side of the array. It is not feasible to constrain the beam to only cover the detectors, as seen in Figure 2.4, since that would mean the outer rows receive less of the dose than their neighboring rows, and this will degrade the quality of reconstructions. That portion of the beam that falls beyond the detectors is called overbeaming.

In practice, the proportion of the total dose from overranging diminishes as the scan length gets longer since overranging at the ends is the same whether the scan length is short or long. This means that overranging can add significantly to total dose on focused exams like temporal bone scans, but its contribution is considered insignificant on studies that cover a lot of anatomy, such as chest-abdomen-pelvis scans. In cases where the dose contribution from overranging is considered substantial, you should consider using axial mode imaging to better constrain the dose. It may be hard to really gauge the impact of overranging, however, since some of the current generation of large-array scanners can limit the added dose from overranging through sophisticated collimation at both ends of the scan range. And,

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(A) region of interest

(B) axial mode

(C) helical mode

Figure 2.6 When covering a short segment of anatomy (A) using axial mode imaging, the scan (B) essentially begins and ends at the top and bottom of the region of interest. When you choose helical mode, where image reconstruction depends on estimation of attenuation between two known points, the scan must begin above the region of interest and end below (C). When using wide arrays and high pitch values, overranging can add substantially to the total dose.

for some scanners with very wide detector arrays, helical imaging is not even necessary for head scans since they can be performed with a single axial rotation that eliminates overranging altogether.

Pitch The choice of pitch is always a compromise among scan speed, patient dose, and image resolution. It is useful to consider using a pitch of 1 as a starting point (Figure 2.7). Pitch 1 means that there are no gaps in the path covered by the X-ray beam, and no tissue is exposed more than once. Increasing the pitch above 1 means that gaps will appear in coverage by the X-ray beam so that some tissue receives much less radiation. Although this would result in complete gaps in imaging when using axial mode scanning, because of the interpolation of data these gaps in the helical scan simply mean that the estimations are less accurate. In principle, this can be used to reduce dose as long as adequate image quality is preserved, and increasing pitch allows faster coverage of the anatomy. For example, this is a good option to consider for CTA of the chest, where resolution demands are less than for brain CTA but rapid coverage of a large section of anatomy is necessary. With the exception of cardiac imaging, it is uncommon to use a pitch of less than 1, but it may be helpful for high-detail exams such as temporal bone scans. With low pitch, there is actually overlap

Radiation Safety and Risks

of the X-ray path over the body. This results in increased dose in those areas, but, this oversampling allows a more accurate measurement of attenuation values and lower noise. Low pitch also requires more time to cover the same anatomy.

Tube Rotation Time Rotation time and mA are interdependent and are sometimes combined in the term milliampere seconds (mAs). That is because the number of X-rays produced is a function of the magnitude of the tube current and the amount of time the tube is turned on. Since the tube is generally on the entire time it is moving, the mAs can be decreased both by using a lower tube current or by spinning the gantry faster so that the rotation is completed faster. On some scanners, however, decreasing the tube rotation time will cause the scanner to automatically increase the tube current to offset the shorter time the tube is on. This results in no net change in mAs and, as a result, there is no dose reduction and little impact on image quality. But the faster rotation speed may prove to be an advantage since, for CTA, it can increase the chances of imaging the contrast during the arterial phase in the brain. But for detailed imaging, like temporal bone exams or routine brain imaging, where there is usually no advantage to decreasing the total scan time from 10 seconds to 5 seconds, it is usually better to use a longer rotation time since it provides better a signal-to-noise ratio (SNR) at equivalent dose.

mA and kV There are many factors that you can alter to lower dose. These include X-ray tube kV and mA, tube rotation time, detector collimation, tube collimation, pitch, reconstructed slice thickness,

Figure 2.7 This illustration shows the wide gaps between the X-ray beam wraps when using a high pitch (upper image) compared with a low pitch (below ). It also shows why so much more anatomic coverage is possible in the same time using the larger pitch.

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reconstruction kernel, and more. The two that are usually considered the most important contributors to total patient dose are the X-ray tube electrical current (mA) and the potential (kV). Voltage and amperage are standard measurements of electrical current that acknowledge the contributions of Andre-Marie Ampere and Alessandro Volta to our understanding of electricity. If you have trouble understanding what these measures of electrical energy mean—and many do—you may fi nd it is easier to picture water instead of electrical current. For example, you can consider that voltage is equivalent to the water pressure at the end of a hose, whereas amperage indicates the actual volume of water flowing from the hose. My grandfather explained this to me when I was child (which I suppose says a lot about my grandfather and my childhood) by using two streams of water emerging from holes in a large water tank. A large hole just below the top of the tank would allow a lot of water to escape but at very low pressure. This is comparable to the low voltage but high amperage current provided by a car battery. Now visualize a very small hole near the bottom of the tank. That hole would allow a small, low-volume stream to shoot out of the tank but under tremendous pressure. In electrical terms, that stream would be equivalent to the high voltage and low amperage current typically used to drive the flash in your camera or phone. Static electricity, which we consider harmless, is measured in thousands of volts but with very low amperage. When considering an X-ray tube, increasing the tube current (referred to in milliamperes or mA) will result in more X-rays created at the anode, but the energy of those X-rays is determined by the electrical potential across the tube, which is measured in kilovolts (kV). A higher kV means a greater electrical potential across the tube, and that leads to higher mean energy X-rays emerging from the anode, with greater potential for tissue penetration. A change in kVp, however, alters both the energy of the X-rays emerging from the tube and their number. Although we frequently talk about CT scans in terms of a single kV value (i.e., kV 120), you should keep in mind that this number represents the peak voltage across the X-ray tube and not the average energy of the X-rays. Although kV reflects the energy range of the X-rays, the beam is composed of X-rays at multiple energies—it is polychromatic. The actual energy of the X-rays in the beam can be described in terms of thousand-electron volts or keV. At a kVp of 80, the lowest X-ray energies in the beam could be as low as 20keV, but no X-rays will be higher than 80keV. The mean energy within this polychromatic X-ray beam is about one-third to one-half of the peak energy predicted by the kVp. The use of specially shaped metal filters between the X-ray source and the patient can substantially alter the energy range of the beam. It is commonplace to use a filter that strips out the very lowest energy X-rays since these contribute to dose without contributing to the image. Another benefit to using a filter is that by narrowing the range of X-ray energies in the beam, it reduces the beam hardening that occurs within the patient. The filter is also shaped to better suit the shape of the patient, to reduce the number of X-rays at the edges where there is usually less tissue to penetrate than at the center. The magnitude of the tube current corresponds roughly to the number of X-rays emerging from the X-ray tube, and the imager should use no more and no less necessary for diagnosis. The dose relationship to mA is linear—lower the mA by half, and you cut the dose by half. But whenever you decrease mA by half, you need to remember that noise increases by about 40%. It is also important to recognize that tube current (mA) may be expressed in several different ways that can lead to some confusion when you are changing scan parameters. Tube current can be expressed as mA, mAs, or effective mAs. That last term incorporates the rotation time and tube current, as well as the scan

Radiation Safety and Risks

pitch. Effective mAs (mAseff) is simply the mAs divided by the pitch. For example, if the pitch is decreased from 1 to 0.5, the mAseff doubles. Why do you need to know this? So that you are not surprised to find that your scanner automatically decreased the tube current when you selected a pitch below l to improve the quality of your temporal bone exams. The scanner software may be set up to keep effective mAs constant, and you will fi nd that the CTDIvol measure of the scan did not change with the alteration of pitch, nor do your scans look better. That is why you need to check the values of mAs and dose for the resulting scans whenever you alter scan parameters. The tube potential or kVp determines both the energy of the X-rays and also the number created at the anode of the X-ray tube. For many users, a kVp of 120 is adequate for nearly all scans, but this “onesize-fits all” approach is quite different from conventional X-ray imaging, in which the kV is frequently adjusted to suit the imaging task. Many believe that CT imaging should take more advantage of this approach since the ideal kV value should be based on the size of the patient, imaging goals, and whether CT contrast is administered, since contrast has a powerful effect on radiation dose to the patient.

When to Increase kV The choice of kV has a much larger impact on patient dose than does the choice of mA. That is because patient radiation dose increases proportionately to mA but increases approximately by the square of kV. For example, an increase of kV from just 120 to 140 results in a 30% increase in patient dose. Although using low kV for CT is desirable when possible, since low-energy X-rays are attenuated more easily, it may be necessary in some circumstances to actually increase kV above 120 (Figures 2.8, 2.9). This is usually the case in large patients or in patients with implanted metal. Although it is reasonable to consider increasing mA first to accommodate requirements when imaging large patients, in some cases the tube limits may be exceeded by the requirements. And it does not seem reasonable to use more low-energy X-rays when what is really required are higher energy X-rays that have better penetration. That is certainly the case when imaging patients with titanium aneurysm clips. Adding more low-energy X-rays in that situation makes little difference toward minimizing artifacts since they are related to photon starvation, and this is best addressed by using X-rays with better penetration. And keep in mind that using a lower kV does not mean there will be a decrease in X-ray dose in all cases. Because noise increases as kV decreases, it is possible that the additional X-rays needed to offset the increased noise may result in a net increase in patient dose. This is more likely to occur in large patients and with non-contrast CT scans. New scanner software that incorporates both kV, mA adjustments based on the scout view, and not just tube current as nearly all AEC does now, should help the user optimization these two factors.

When to Decrease kV Adequate low-kV CT imaging is frequently possible when performing body imaging in children and thin adults. This is because the quality of penetration of high-kV X-rays may not be necessary for that patient population. Although decreasing kV is accompanied by an increase in image noise and

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Figure 2.8–2.9 On these sagittal reconstructions from a neck CT demonstrate image quality is worse in the lower cervical-upper thoracic segments. This is because insufficient X-rays were arriving at the detectors at those levels because of the patient’s large shoulders. Although increased mA is a consideration, in such cases, increasing kV may be necessary to provide adequate imaging and there is an upper limit to tube mA.

beam hardening artifacts, the improvement in contrast conspicuity whenever using iodinated contrast in most circumstances offsets these disadvantages. You should recall that iodine is more evident on low-kV CT scans because of the photoelectric interaction of low-energy X-rays with iodine. Since X-ray beams are polychromatic, the X-ray beam generated using kV 80 will contain many X-rays close to the k-edge of iodine, around 30keV. The increased noise at low kV can be made less evident with an increase in mA while still providing a lower total dose. Although brain imaging is nearly always performed with kV 120, for high-dose exams like brain perfusion, this property of iodine allows adequate imaging using kV 80 (Figures 2.10, 2.11). Although not commonly considered in practice, the increased conspicuity of iodine should also allow the use of less contrast for some applications. This could prove to be an advantage for patients with marginal renal function.

Detector Collimation and Slice Reconstruction It is preferable to use narrow detector collimation to improve resolution and decrease artifacts, as long as sufficient dose is utilized. Since there are fewer photons collected per detector when using detector collimation of less than 1mm compared with say 2mm detector collimation the SNR reflected on images will be lower. Noise can be estimated as 1/[square root of slice thickness]. This means that a 10mm slice has three times less noise than a 1mm slice.

Radiation Safety and Risks

Figures 2.10–2.11 These images of the same level in the brain were created simultaneously on a dual-energy scanner. You will notice that, in Figure 2.10 (left), the lower kV image, the contrast staining in the patient’s left hemisphere is much more apparent compared with Figure 2.11 (right), the kV140 image. This improvement in contrast conspicuity explains why low kV CT imaging should be considered for abdominal and pelvic scans since the decrease in signal-to-noise ratio may be offset by the improvement in contrast and it substantially reduces patient dose.

In practice, the primary use of thin detector collimation proves to be less of a problem than one might expect. That is because the data from each thin detector can be combined with three or more neighbors to provide image reconstructions at 5mm, with considerably more SNR than if images were reconstructed at a displayed slice width of 0.625mm. However, whenever the very thin sections are used primarily for diagnosis, the mAs will need to be increased as detector collimation decreases to preserve image quality. A variation of using thin detectors and thick reconstructions that preserves SNR and some of the benefits of viewing the thin sections directly is to reconstruct relatively thick images but at a smaller increment than the displayed slice thickness. For example, you have the option of generating 2mm reconstructions but at 0.5 mm intervals. The only real disadvantage of this approach is that it increases the total number of slices necessary for storage and review. You might wonder: “If I am going to combine data from detector rows, and SNR is better with thicker detector collimation, why use thin detector collimation at all?” First, keep in mind that if you perform the scan using thick detector collimation you cannot later go back later and reconstruct thinner sections. There are other benefits of thin detector collimation, however. It decreases the artifacts from partial volume that may appear as indistinct edges of structures on helical reconstructions and volume averaging. It also reduces the beam hardening artifacts that degrade posterior fossa imaging on brain CT.

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Filter or Kernel The terms filter and kernel are used interchangeably and refer to the method of raw data reconstruction after the scan is acquired. There is potentially less confusion if you use the word “fi lter” to refer to the metal plate that is designed to optimize the X-ray beam and use “kernel” to refer to your choice of image reconstruction algorithm. The choice of kernel proves to be significant in any dose reduction project since apparent image noise will influence decisions regarding the dose necessary for adequate scanning. So, although the reconstruction technique is not commonly considered a dose reduction tool, you should at least be aware of how this parameter may influence your perception of image quality.

Automatic Exposure Control It is one thing to select ideal scan parameters when examining a small part of the anatomy, such as the temporal bones, but quite another when performing a scan that includes several body parts of very different thickness, shape, and composition, such as chest and abdomen. The problem with these parts of the anatomy that have wide variations in shape and thickness is that the ideal choice of kV and mA for one portion of the exam will prove to be either too much or too little for another portion covered during the same scan. For example, the appropriate choice for tube current in the chest will be insufficient for imaging the abdomen. Even within a single slice through the chest, what may be an appropriate dose in the anterior to posterior (AP) direction may prove to be insufficient to penetrate side to side across both shoulders. This has become a more commonplace problem as helical imaging has allowed coverage of long sections of the body in a single scan. To better allow the scanner to match the tube current to the requirements of varying anatomy, most manufacturers offer some variation of automatic exposure control (AEC) on their scanners. This dose reduction tool is based on one or two scanograms that are created at the start of the exam (Figure 2.12). Automatic exposure control usually modifies only the tube current (i.e., mA) during the scan, not the kV. That is why it is sometimes called tube current modulation. By modifying the dose during the scan to match the attenuation of that part of the body, both from slice to slice and within the same slice, significant dose reduction is possible without image degradation. Automatic exposure control should be considered for CT imaging whenever the scan covers a portion of the body where there are significant variations in thickness or attenuation. On some scanners continuous adjustments are made even during a single rotation. This technique is called angular adjustment and it provides more X-rays during the part of the rotation from side to side through the shoulders than when imaging from front to back. The use of that software tool can provide a lower overall dose with better quality, assuming the reference values are correctly set. There are four important things you need to understand about AEC to use it properly. First, as suggested by the word “automatic,” the user needs to recognize that they have given up a small degree of control over the scan parameters and therefore the fi nal patient dose. Second, you need to understand that AEC does not “automatically” lower dose. All it does is match tube current to the anatomy based on some target value, such as noise level or mAs equivalent.

Tube current

Radiation Safety and Risks

Z-axis position

Figure 2.12 This illustration shows how the scanner uses the scanogram obtained at the start of the exam to calculate a dose profile for the patient. This profile is then used to automatically make adjustments to the tube current (mA) during the scan acquisition. These adjustments are also based on some predetermined imaging benchmark of quality e.g. image noise or a preselected mA equivalent. Even though intended as a dose reduction tool, if the desired benchmark is set high, the scanner may use a dose that is higher than you might have selected without AEC.

It is up to the user to pick an appropriate mA or noise value in order to provide acceptable images but at the lowest possible dose. Third, when imaging thin patients and children, is very important to decide on the appropriate tube voltage fi rst. Automatic exposure control on nearly all scanners will only modify tube current. For thin patients and children, fi rst consider decreasing kV for body imaging since that has a more substantial impact on dose than does AEC. When modifying kV, keep in mind that AEC quality settings are commonly relevant for only one kV selection. For example, if you choose to decrease tube voltage and also activate AEC, the increase in noise at the lower kV may lead to the AEC to apply a much higher mA value than necessary. Fourth, keep in mind that AEC is intended to accommodate changing anatomy. One instance where AEC may result in excessive patient dose is when it is used during continuous imaging such as CT brain perfusion. In one hospital, it was reported that 200 patients received unusually high doses of radiation during CT perfusion exams in part because AEC was turned on for the brain perfusion studies. This was selected with the intent to reduce dose, but the software “automatically” provided a very high dose. The images were very likely of excellent quality, which may explain why this problem went undetected for so long.

Shielding Since X-ray dose at diagnostic energies is much higher in superficial tissues, there has been some interest in using radiation shields over particularly sensitive tissues like breast, thyroid, and eyes. These shields are usually made of bismuth, which has desirable attenuation characteristics compared with lead, for example (i.e., X-ray attenuation without significant metal artifact). These shields are not in widespread use in part because they add cost when used for eyes (eye shields are one-time use), and they add to exam time elsewhere in the body since they need to be applied after the scanogram but before the scan. This is important to keep in mind whenever AEC is used with shields since it is the scanogram that is used to select the tube current on most scanners. If the shield is put in

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place prior to the scanogram, say at the beginning of the scan, the software would direct the AEC to increase tube current at the level of the shield and that would effectively offset any benefit to shielding. In spite of these current limitations, shielding should at least be considered as part of any comprehensive dose reduction program.

Iterative Reconstruction The first CT scans used a purely mathematical reconstruction of the collected data to create images. That was found to be too slow for clinical imaging and all current scanners routinely use some variety of back-projection. With increasing concern about patient dose from CT and the decreasing cost of computer hardware, it has become commonplace for vendors to offer some variety of mathematical reconstruction called iterative reconstruction. By referring back to the source data, and in some versions accounting for the fixed noise in the system, iterative reconstruction can improve the imaging SNR and that can be used for better images at the same dose or equivalent imaging at a lower dose. Although there is no question it can help with dose reduction, its impact on diagnosis is still uncertain. With any new reconstruction tool, it would be expected that new artifacts may appear, and altered image contrast may alter sensitivity. Early reports, however, indicate that dose reductions of at least 50% for body CT imaging and 10–20% for brain imaging are to be expected once iterative reconstruction is fully integrated into clinical scanners. It seems highly likely that some version of pure mathematical reconstruction or a blend of back-projection and iterative reconstruction will become routine for CT reconstructions.

Pediatric CT Imaging Children are considered most vulnerable to the stochastic effects of radiation because any DNA damage that occurs is amplified by both the number of years of life and by the fact that it occurs before the reproductive years. That is why the effective dose for a head scan in a child is higher than for a head scan with the same DLP in an adult. Another factor to consider is that CT is being used more frequently to diagnosis nonmalignant diseases, such as appendicitis in young patients, and in nearly all cases of head trauma. In fact, the likelihood of getting some sort of CT imaging after any sort of trauma is quite high at many U.S. hospitals since clinical signs may be misleading. This broadening of indications for CT and its widespread availability have led to an explosion of utilization in spite of increased access to MR during the same time period. So, although it is important to minimize X-ray dose for all CT patients, the benefit of reducing dose is even larger for children. In response to this concern, considerable attention to dose reduction has already occurred at most pediatric centers. One common method to reduce patient dose in children is to decrease the kV of the exam since there may be no need for the same X-ray penetration in children as in adults. For body imaging with contrast agents there can be substantial benefits to imaging with a lower kV. Because dose increases as the square of kV, decreasing kVp from 120 to 80kVp results in a 65% decrease in

Radiation Safety and Risks

dose at the same mA. Although noise increases, the improved image contrast because of the photoelectric interaction of low-energy X-rays with iodine provides acceptable image quality after an adjustment of mA. Some useful CT scan protocol recommendations can be found on the web by searching for the term “Image Gently.” Although you may hope to fi nd a set of protocols based on age and size, remember that because of the very complex interplay between X-ray tubes, fi lters, detectors, and scanner dimensions, it is unlikely that what is optimal for one site will be optimal for another. In an attempt to accommodate these differences in scanners, the recommendations you will fi nd on that website are expressed in terms of percent change from adult technique instead of absolute values. Iterative image reconstruction will very likely offer opportunities to substantially reduce dose for both body and head CT imaging in children, and so you should be aware of this option and consider using it once it becomes available on your scanner. SUGGESTED READINGS Nievelstein RAJ, van Dam IM, van der Molen AJ. Multidetector CT in children: Current concepts and dose reduction strategies. Pediatr Radiol. 2010;40:1324–1344. Lee CH, Goo JM, Lee HF, Joon S, Park CM, Chun EF, Im JG. Radiation dose modulation techniques in the multidetector CT era: From basics to practice. Radiographics. 2008;28:1451–1459. Golding SJ. Radiation exposure in CT: What is the professionally responsible approach? Radiology. 2010;255:683–686. Kalra MK, Maher MM, Toth TL, Hambert LM, Blake MA, Shepard J, Saini S. Strategies of CT radiation dose optimization. Radiology. 2004;230:619–628. Verdun FR, Bochud F, Gudinchet F, Aroua A, Schynder P, Meuli R. Radiation risk: What you should know to tell your patient. RadioGraphics. 2008;28:1807–1816. Schilham A, Molen A, Prokop M, Jong HW. Overranging at multisection CT: An underestimated source of excess radiation exposure. RadioGraphics. 2010;30:1057–1067. Tamm EP, Rong JX, Cody DD, Ernst RD, Fitzgerald NE, Kundra V. Quality initiatives. CT radiation dose reduction: How to implement change without sacrificing diagnostic quality. RadioGraphics. 2011;31:1823–1832. Huda W, Ogden KM, Khorasani MR. Converting dose-length product to effective dose at CT. Radiology. 2008;248:3.

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CARDIAC CT IMAGING TECHNIQUES Supratik Moulik and Harold Litt

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Temporal Resolution: The Key to Cardiac Imaging The basic principle of cardiac imaging is to acquire the scan data during a motion-free portion of the cardiac cycle (Figure 3.1). This is no small thing because of the continuous, rhythmic motion of the heart. The solution is to identify a relatively motion-free part of the cardiac cycle (i.e., end diastole) and acquire images during that short window in time of approximately 120–170 msec. It is also possible, with retrospective gating, to dissect from a helical computed tomography (CT) scan images of the heart at multiple time points in the cardiac cycle, but this partitioning of data is limited by the temporal resolution. Temporal resolution is an important concept toward your understanding of motion-free imaging of the heart. In the language of cardiac CT, temporal resolution is the shortest time necessary to acquire a detector array’s worth of information sufficient to reconstruct images (Figure 3.2). This time varies with the design and capabilities of the scanner, of course, but gantry rotation time is the most important factor to consider. If we are to provide the best temporal resolution, gantry rotation time needs to be as short as possible. But there are physical limitations on how fast the gantry can rotate since tremendous centrifugal forces are generated at high rotation speeds because of the combined weight of the X-ray tube, filters, collimators, and detector apparatus, which includes the septa and the detectors themselves (Figure 3.3). The magnitude of these forces can be calculated using the angular velocity (ω) of the gantry, the radius of the gantry track, and the total mass of the spinning hardware. If you have ever had an unbalanced washing machine stop during a spin cycle when you wash a heavy blanket, you know intuitively that the heavier the spinning object and the faster the spin, the greater the force. Since multidetector scanners require heavy X-ray tubes to allow continuous imaging, and the scanner diameter needs to be large enough to fit patients, there is a limit to gantry rotation time since the forces increase as the

Figure 3.1. Normal cardiac cycle.

5 segment

171 msec

10 segment

86 msec*

20 segment

43 msec

Figure 3.2 The temporal resolution determines how finely the information can be divided within a given cardiac cycle. For example, when there is high temporal resolution, 20 segments can be created compared with only 5 at lower temporal resolution.

Cardiac CT Imaging Techniques

Figure 3.3 Single-source CT scanner with third generation geometry.

Table 3.1 Gantry rotation speed for various generations of CT scanners

Gantry rotation time First generation

5 min

Second and third generation (narrow and wide fan beam)

2 sec

Electron beam

100 msec

Fan beam MDCT

0.3–1 second MDCT

Cone beam MDCT

270–500 msec

Dual-source

285–332 msec

square of angular velocity. Although flat-panel detectors have been developed that are significantly lighter than current multidetector arrays, the need to balance the weight of the X-ray tube largely negates any potential weight benefit of flat-panel detectors (remember the washing machine). The rotation speed and temporal resolution of the various generation scanners are provided in Table 3.1.

How Temporal Resolution Can Be Shorter than Gantry Rotation Time Given the physical limitations on gantry rotation speed, further improvement in temporal resolution can also be achieved with advanced reconstruction methods. The two most widely utilized methods are partial scan acquisition and segmented reconstruction. Each of these techniques has its

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Figure 3.4 This illustration shows how the complete 360-degree dataset for CT reconstruction can be obtained in thirds over three consecutive cardiac cycles.

advantages and disadvantages, although both provide significant improvements in temporal resolution (Figure 3.4). SEGMENTED ACQUISITION Segmented acquisition means that the data necessary to reconstruct an image are acquired in pieces during multiple cardiac cycles. For nearly all multidetector scanners, it requires a half a second or more for one complete gantry rotation. But because the optimal window for motion-free cardiac imaging—end diastolic imaging—is only about 150 msec long, this means that, during that single gantry rotation, the heart would be viewed in multiple phases of the cardiac cycle. By using data from just one-third of the full rotation, however, and collecting the rest of the necessary data over the following two cardiac cycles, temporal resolution can be shorter than the gantry rotation time, but only if the gantry rotation is decoupled (phase offset) from the heart rate. For example, if the gantry speed is one rotation per second and the heart rate is 60 beats per minute (bpm), then each gantry rotation would result in an identical dataset that will be insufficient for reconstruction. That occurs, of course, because a full rotation is necessary for reconstruction, and, although that can be pieced together from multiple rotations, each fraction of the 360-degree rotation must occur in the correct temporal window with respect to the cardiac cycle. You can think of it in this way: if you want to see all the people on a merry-go-round that completes one rotation every minute, and you are only allowed to look at it for intervals of 10 seconds, then you must vary the intervals of your observations. If you were to look up every minute on the minute, you would see the same people each time. This issue is most commonly resolved by allowing the scanner to perform variable pitch scanning, such that the pitch is continuously modulated based on the heart rate. This entire process of spreading the acquisition of image data over multiple cardiac cycles is known as segmented acquisition. In the ideal case, when the gantry rotation and heart rate are completely decoupled and there is no overlap in projection data, the improvement in temporal resolution can be expressed as gantry

Cardiac CT Imaging Techniques

rotation time divided by the number cycles used for acquisition. For example, if you were to use four rotations for data acquisition, and gantry rotation time was 1 second, then the temporal resolution will be only 125 msec. Imperfect decoupling and other practical factors tend to degrade the actual improvement in temporal resolution, however. The radiation dose to the patient is also higher with this method. Ideally, when using segmented acquisition, there should be identical motion and position of the heart from one cardiac cycle to the next. This would imply a consistent breath-hold, with no significant variability in the heart rate or cardiac contractility. In practice, since most patients have some degree of beat-to-beat variability in heart rate, as well as imperfect breath-holding, segmented acquisition may have limited utility compared to other techniques. PARTIAL SCAN When considering the projection data acquired from a scanner with narrow beam geometry (Figure 3.5), it would seem that the data acquired from the horizontal tube position should be identical to the data acquired when the source and detector are reversed, if we discount the differences due to the diverging beam. This observation can be exploited to decrease scan acquisition time by performing the image reconstruction from only 180 degrees of data compared with the usual 360-degree dataset. Using this approach, the CT scan temporal resolution can be improved by roughly a factor of 2 without changing the gantry rotation speed at all. DUAL-SOURCE CT Further improvement in scanner temporal resolution for cardiac CT has been made possible through the introduction of multisource scanners. The most widely used variation for cardiac CT scanning positions the two source-detector pairs at approximately 90 degrees with respect to each other (Figure 3.6). The

Figure 3.5 Justification for half-scan technique. The measured projections should be identical with beam at zero and 180 degrees.

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Figure 3.6 This view down the bore of dual-source CT scanner shows the two tubes at 90 degrees in the gantry, here shown at the top and right side of the bore. Their corresponding detectors are evident opposite each tube.

actual gantry rotation speed for one commercially available unit is approximately one-third of a second, which is remarkable considering the mass of the spinning hardware in such a design. If the scan data from the two tube-detector pairs is combined to create a single image, then the gantry needs only to rotate 90 degrees in order to obtain sufficient data to reconstruct a respectable image. When utilizing such a scanner using partial scan technique, only one-quarter of a gantry rotation in necessary with a dual-source scanner. Using this combination of acquisition technique and scanner hardware, the temporal resolution is 83 msec. And, when using both segmented acquisition and the partial scan technique with a dualsource scanner, one can expect temporal resolution that is actually superior to electron beam CT while still providing motion-free imaging within the range of normal heart rates (60–100 bpm).

Electrocardiogram Synchronization Motion-free CT cardiac imaging is necessary so that the CT scan data acquisition can be synchronized to the patient’s electrocardiogram (ECG). Two widely implemented techniques for linking these two data streams are called prospective triggering and retrospective gating. It is important to consider when to acquire the images during the cardiac cycle since an increase in heart rate disproportionally shortens the diastolic fi lling period, while the systolic phase will remain relatively constant over the common range of heart rates. For low heart rates (3) scanning is a useful method for dealing with patient motion since it allows performance of an entire chest or abdominal CT in less than 3 seconds. Compared to a retrospectively gated CTA acquired at a pitch of 0.3, these highpitch techniques can achieve relatively motion-free scanning in one-tenth the time. Finally, positioning aids such as wedges and pillows, which allow the patient to rest comfortably in the desired position, may be of value. Figure 4A.2.4 illustrates how patient motion can be tracked via markers. The image (A) represents the patient in a neutral position. The marker, which is positioned above the patient, is at the neutral position as well. The second image (B) represents patient motion in the y-direction, which is indicated by a shift in the marker position. The fi nal image (C) is a return to the neutral position. Modern CT scanners offer some other techniques to compensate for motion artifact; these include overscanning and software correction algorithms during reconstruction. Overscanning means that more than a 360-degree gantry rotation is used and the data from those additional degrees of coverage are then averaged with the initial portions of the scan to minimize the motion artifacts during any given slice. Other motion compensation methods have been proposed that combine multipoint, skin marker motion tracking in a manner similar to that used by the movie and video game industries.

Cardiac CT Artifacts and Pitfalls

Figure 4A.2.2 This sagittal reconstruction from the scan seen in Figure 4A.2.1 demonstrates a stair-step artifact due to patient motion during the cervico-thoracic portion of the scan acquisition. Nevertheless, the lower part of the study was still of diagnostic quality.

The underlying premise is to treat the body as a rigid structure and to track its motion in three dimensions. This motion data can be then be added to the projection data at corresponding time points. The limitation of this technique is that the human body is not rigid, and patients do not move stiffly on the CT table. Although it seems likely that further developments in optical motion capture techniques will translate into improved motion suppression, no such systems are currently in use.

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Figure 4A.2.3 In this graphical representation of motion artifact during an axial acquisition, the top image shows the expected relationships without motion, whereas the bottom image shows how one slice appears offset relative to its neighbor due to motion.

Figure 4A.2.4 These images illustrate how patient motion can be tracked via markers. The left image represents the patient in a neutral position. The marker, which is positioned above the patient, is at the neutral position as well. The second image (middle) represents patient motion in the y-direction, which is indicated by a shift in the marker position. The final image (right) is a return to the neutral position.

Correct answer: 3, 2, 4, 1

Cardiac CT Artifacts and Pitfalls

Artifact 3 This patient is a 60-year-old female who presented to the emergency room complaining of acute onset of shortness of breath 3 hours previously (Figure 4A.3.1). She recently returned from vacationing in Europe and also complains of left leg pain and swelling. The patient is alert and oriented ×4, although she remains hypoxic (90% oxygen saturation) and short of breath on 2 L oxygen via nasal cannula. She was unable to maintain the breath-hold during CT pulmonary angiogram performed on a dual-source CT scanner. What is the most reliable way to differentiate respiratory motion from cardiac motion on a chest CT? (1) Look for blurring of cardiac contours on axial images. (2) Look for blurring of pulmonary vascularity on axial images. (3) Look for stepwise artifact in the sternum on sagittal images. (4) Compare pre- and postcontrast images.

Figure 4A.3.1 This axial image from a CT pulmonary angiogram demonstrates respiratory motion artifact that limits the visualization of many pulmonary artery branches.

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(B)

Figure 4A.3.2 Sagittal image (B) in a different patient demonstrates the typical irregularity of the sternum seen with respiratory motion. Lateral scanogram image from the same scan (A) demonstrates a normal sternum in that area. The respiratory motion artifact is seen on an axial image from that level (C).

On cardiac gated studies, differentiating cardiac from respiratory motion helps to separate those cases with cardiac motion that will benefit from postprocessing techniques from those with respiratory motion in which postprocessing will prove to be of limited utility. Blurring of the cardiac contours and pulmonary vascularity on axial images may be evident as a result of both respiratory and cardiac motion (Figure 4A.3.2 A–C). The key to differentiating the two is the presence or absence of chest wall motion artifact, and that will be most evident on the sagittal image of the sternum.

Cardiac CT Artifacts and Pitfalls (C)

Figure 4A.3.2 (Continued)

In general, respiratory motion artifact is deleterious on only a limited portion of a scan, leaving the other parts of the scan unaffected. Similar to gross patient motion discussed previously (Artifact 2), prevention is the most reliable method for dealing with this particular artifact. In the case of respiratory motion artifact, the most important factor to consider is the patient’s respiratory status and factors such as severe chronic obstructive pulmonary disease (COPD) or fibrotic lung disease that limits his or her ability to perform a 10–20 second breath-hold. The most commonly used methods for prevention of respiratory motion artifact include longitudinal axis collimation (i.e., minimizing the volume scanned), caudal-cranial scanning, and high-pitch scanning. Minimizing scan length decreases the time required for image acquisition and breath-hold, as well as minimizing the radiation dose, which is directly proportional to the coverage area. Worse yet, as the breath-hold gets longer, there will be some degree of reflex tachycardia, which exacerbates the motion artifact. For severely hypoxic patients, shortening the scan to allow coverage of the region of interest in a single breath-hold may help to obtain a diagnostic scan that might otherwise be impossible. Caudalcranial scanning can be beneficial in patients being scanned for evaluation of pulmonary embolism since most pulmonary emboli occur in the lower and middle lobe branches; so, by scanning this area fi rst, image quality improves in the region of greatest interest. High-pitch (3.4) scanning is a new method available on some scanners that allows for triggered, full-volume coverage with nonoverlapping helical acquisition that utilizes two X-ray sources, as well as specialized reconstruction algorithms.

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Figure 4A.3.3 Axial image through lung bases.

Figure 4A.3.4 Four-dimensional CT. Images from a 4D CT were obtained for radiation therapy planning in a patient with right lower lobe lung cancer. The CT acquisition was performed continuously throughout the respiratory cycle at a low pitch to allow reconstruction of image sets at eight different phases throughout the respiratory cycle. Each image shown is a two-dimensional (2D) projection created from the retrospective phase dataset (12% on the left). The arrows mark the position of the right sixth and eighth ribs; the position of the right lower lobe mass varies throughout the respiratory cycle relative to these ribs.

This axial image through the lung bases (Figure 4A.3.3) from a different patient illustrates the difficulty in evaluating adequate images when the question is pulmonary embolism, as respiratory motion mimics vascular fi lling defects. Other more complicated and uncommonly used methods for respiratory motion compensation exist and provide varying degrees of utility. In the setting of radiotherapy/proton beam therapy for treatment of lung tumors, four-dimensional (4D) CT scanning is being investigated for potential clinical applications (Figure 4A.3.4).

Correct answer: 3

Cardiac CT Artifacts and Pitfalls

Artifact 4 This patient is a 55-year-old female with a history of intermittent chest pain who presented to the emergency room with worsening shortness of breath (Figure 4A.4.1 A,B). The patient had negative cardiac enzymes, and a retrospectively gated cardiac CTA was performed. During the image acquisition, the patient experienced reflex tachycardia up to 80 beats per minute (bpm) from a baseline rate of 60 bpm. Which of the following is correct regarding tube current modulation during cardiac CTA: (1) Irregular heart rates allow using a lower radiation dose. (2) The method used for predicting the appropriate dose modulation is similar to that used in prospectively triggered acquisitions. (3) Increased heart rates facilitate dose reduction. (4) Image quality will be the same in all phases of the cardiac cycle.

(A)

(B)

Figure 4A.4.1 Volume-rendered (A) and oblique coronal (B) images from a retrospectively gated coronary CT angiogram (CTA) demonstrate bands of high image noise artifact resulting from improper tube current modulation.

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Figure 4A.4.2 Electrocardiogram tracing from a retrospectively gated cardiac CT. The dark bars represent the 75% phase set as end diastole. The shaded areas throughout the rest of the cardiac cycle represent the lower tube current portions of the acquisition.

With tube current modulation, the scanner tries to predict the timing of the selected end diastolic phase based on the previous 3–5 cardiac cycles, a method very similar to prospectively triggered cardiac CT studies. Irregular heart rates make it difficult or impossible to predict the appropriate timing of the next cardiac cycle reliably, which increases the portion of the cardiac cycle requiring the full tube current. Thus, answer 1 is incorrect. Increased heart rate decreases the effectiveness of tube current modulation because the overall cardiac cycle is shorter, which decreases the time window when current can be lowered. Thus, answer 3 is incorrect. The image quality for any given slice depends largely on the quantity of the X-rays that contribute to the image, and that is determined by the tube current. During the phases of the cardiac cycle when the tube current is decreased for dose reduction, there will be more noise compared with the phases when the full tube current is used. Thus, answer 4 is incorrect. Tube current modulation is a widely used dose reduction technique on modern CT scanners. The radiation dose is then optimized based on the specific patient’s density profi le and cardiac phase without significant image quality penalty. In cardiac CT acquisition, tube current modulation provides the benefits of prospective triggering to be partially realized in retrospectively gated studies. The basic principle is that the relatively motion-free parts of the cardiac cycle are end systole (usually occurring at 35% of the R-R interval) and end diastole (at 75%). The projection data for the remainder of the cardiac cycle are used primarily for functional analysis since that does not require the same image quality or spatial resolution as a coronary artery evaluation. Similar to the techniques utilized for prospective triggering, the tube current can be decreased to a fraction (4–20%) of the full dose during phases of the cardiac cycle that are unlikely to provide motion-free imaging of the coronary arteries. The fraction of the cardiac cycle that receives the reduced dose is dependent primarily on the rate and regularity of the heartbeat. For low heart rates, the end diastolic window for image acquisition is relatively short compared to the length of the overall cardiac cycle (e.g., 150 msec of a 1,000 msec cycle is 15%), whereas rapid rates mean an increase in this percentage (e.g., 125 msec of a 600 msec cycle is 21%). Therefore, the dose reduction afforded by tube current modulation is dependent not only on the variability of the heart rate, but also on the heart rate itself (Figure 4A.4.2).

Cardiac CT Artifacts and Pitfalls (A)

(B)

Figure 4A.4.3 Short axis images from a retrospectively gated coronary CT angiogram (CTA) reconstructed at 75% (A) and 95% (B) of the R-R interval with tube current modulation active.

In the setting of normal sinus rhythm with a low variability from beat to beat, tube current modulation can accurately predict subsequent cycles and adjust the tube current appropriately (Figure 4A.4.3 A,B). With a variable heart rate, however, it is more difficult to predict the timing of the next cardiac cycle, so the scanner must widen the window of full tube current to ensure that the true end diastolic phase is covered. As a result, in patients with a high degree of heart rate variability or atrial fibrillation, tube current modulation will provide very little dose reduction. Since projection data are acquired continuously throughout the cardiac cycle, it is possible to reconstruct images during any cardiac phase. As demonstrated in this case, when tube current modulation is active, the image noise increases in the dose modulated portions of the scan, which results in a grainy appearance (Figure 4A.4.4). However, those images are still sufficient for functional

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evaluation, although their poorer SNR may make them inadequate for evaluation of the smaller (10 HU) due to partial volume averaging or whenever the cyst is fi lled with bloody or proteinaceous fluid (called a complicated cyst). Partial volume averaging artifact can be eliminated as a cause by viewing thin slice reconstructions. The diagnosis of a complicated cyst can be established whenever an otherwise benign-appearing cyst has high attenuation on a noncontrast scan and there is no change in the fluid attenuation after IV administration of contrast material. Imaging research using phantoms has shown that renal cyst pseudo-enhancement is independent of partial volume averaging and that the cyst attenuation increases with more detector rows and the use of a higher tube potential (e.g., the use of 140 KVp compared to 90 KVp). Other independent factors predisposing to renal cyst pseudo-enhancement include intrarenal location, small size of the lesion, and imaging during peak renal parenchymal enhancement. The best theory to date for the cause of pseudo-enhancement is that it is the result of the combination of a beam hardening correction and the helical image reconstruction algorithm used by many manufacturers. Remember that beam hardening does not require bone or other high-attenuation tissues to occur. CT imaging of a uniform fluid phantom will appear to have a central low-attenuation zone to the miscalculation of the true attenuation in the middle due to beam hardening. This artifact, at least in a uniform phantom, can be corrected by using a correction factor during the reconstruction to arbitrarily increase attenuation values in the center but this correction has the potential to create artifacts elsewhere.

Correct answer: 4

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Pitfall 2 There is a round area of low attenuation in the medial aspect of the pancreatic head (Figure P8.2.1, arrow) in this patient who was being staged for renal cell carcinoma of the left kidney. This is: (1) an enlarged peripancreatic lymph node. (2) a solid neoplasm in the pancreatic head. (3) a cystic lesion in the pancreatic head. (4) normal low attenuation of the nonenhanced superior mesenteric vein (SMV).

Figure P8.2.1 Arterial phase abdominal CT.

Body CT Pitfalls

This is a commonplace fi nding on multiphase imaging of the upper abdomen that is most often performed to improve visualization of the liver or pancreas. This normal structure becomes more conspicuous against the surrounding pancreatic parenchyma when imaging occurs during the arterial phase of enhancement. When imaged later, during the portal venous phase of enhancement (Figure P8.2.2), it is more easily recognized as the SMV. Contrast is exaggerated again since the pancreatic parenchymal enhancement is past peak. Radiologists have long been aware that certain vascular lesions in the liver or pancreas, such as hepatocellular carcinoma, may only be seen during arterial phase CT scans. Multiphase imaging is also used commonly for the evaluation of the kidneys, where a delayed, excretory phase is necessary for opacification of the pyelocalyceal systems and ureters (Figures P8.2.3 A,B and P8.2.4).

Figure P8.2.2 Portal venous phase image, same slice as Figure P8.2.1.

219

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CT IMAGING (A)

(B)

Figure P8.2.3 This patient had a triple-phase CT to characterize a hepatic lesion and on this image again note the low attenuation of the superior mesenteric vein (SMV) on arterial phase imaging (A, arrow ) that later enhances during the portal venous phase (B, arrow ).

Figure P8.2.4 It is important to direct attention to all phases of the imaging. In this patient with cirrhosis and portal hypertension, the low attenuation in the region of the superior mesenteric vein (SMV) is due to nonocclusive thrombus in the SMV (arrow ).

Correct answer: 4

Body CT Pitfalls

Pitfall 3 There appears to be a central low-attenuation fi lling within the deep venous structures extending from the groin to the mid inferior vena cava in this patient with ulcerative colitis (Figures P8.3.1, P8.3.2). The CT scan was performed following IV injection of 100 mL contrast material at a rate of 2 cc/sec, with a scan delay time of 75 seconds after the initiation of the contrast injection. This fi nding indicates: (1) a potentially life-threatening venous thrombosis; further investigation and treatment is warranted. (2) a flow artifact due to insufficient volume of contrast or early scanning after start of injection. (3) the usual appearance of veins. (4) contrast mixing with unopacified venous blood.

Figure P8.3.1 Abdominal CT in a patient with ulcerative colitis.

Figure P8.3.2 Pelvic CT image from same patient seen in Figure P8.3.1.

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This fi nding is most consistent with extensive, nonocclusive venous thrombus, and that puts the patient at risk for a pulmonary embolism. Distinguishing flow artifacts within the veins from nonocclusive thrombus is very important and can be challenging in practice. The venous luminal filling defects in this case are well defi ned and, on axial images, appear as round or polygonal filling defects. In addition, the right external iliac vein is expanded (Figure P8.3.2, arrow), a fi nding that can be seen with acute thrombosis but not with a flow artifact. The goal for imagers is to avoid false-positive diagnoses of venous thrombosis. The best technique for making this distinction between venous thrombus and flow artifact is a delayed post contrast image rather than giving a larger amount of contrast material. In some cases, a venous ultrasound study will be necessary for the timely confirmation or exclusion of deep venous thrombosis (Figure P8.3.3 A,B). (A)

(B)

Figure P8.3.3 These companion cases demonstrate venous flow artifact. This axial CT shows faint low attenuation within the lumen of both common femoral veins (A, arrows) due to incomplete mixing of blood and contrast material. A CT section from another patient who was evaluated for a ventral hernia with CT shows a faint, linear filling defect in the lumen of the left external iliac vein (B, arrow ) typical of a flow artifact.

Correct answer: 1

Body CT Pitfalls

Pitfall 4 On this patient’s abdominal CT with IV contrast enhancement, a small, low-attenuation lesion is evident in the left lower kidney. On 5 mm axial imaging, it has an attenuation value of 42 HU, but of only 14 HU on 2 mm coronal imaging. The difference in apparent attenuation of the cyst contents is due to: (1) beam hardening. (2) pseudo-enhancement. (3) volume averaging. (4) motion.

Figure P8.4.1 Axial CT section with IV contrast enhancement with cyst contents measurement 42HU.

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Figure P8.4.2 Coronal reconstruction, CT of same patient as in Figure P8.4.1. On this view the cyst attenuation measures 14HU.

Volume averaging on CT occurs because the appearance of a voxel represents the average attenuation value of all the different tissues within the voxels (i.e., gas, fluid, fat, bone, etc.). The larger the voxel, the more tissues within, and that may lead to inconsistencies or frank errors in CT representation of attenuation values. The reason the cyst has a higher attenuation value on the axial view is that the slice is thicker and that means the voxel is larger. Since the cyst is in fact very low attenuation—in fact, close to 0 HU, since the voxels that show it include both cyst fluid and the nearby higher-attenuation enhanced kidney—it is displayed with a higher than expected attenuation value. Thin-slice acquisition and reconstruction may provide a more accurate attenuation measurement, as long as the dose is sufficient for an adequate signal-to-noise ratio on these thinner sections.

Correct answer: 3

9

TEST QUESTIONS Alexander C. Mamourian

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Question 1 The CT measure of how much of the total radiation dose contributes to the images is called dose efficiency. It incorporates both the scanner detector and geometric efficiency. Assuming the detector efficiency is the same for all, place these different abdominal imaging techniques in order of dose efficiency from highest to lowest. (1) Axial mode using a 16-row detector array (2) Axial mode using a single-slice scanner (3) Helical mode using a four-row detector array

Question 2 Which of these scan techniques would deliver a lower patient dose for an abdominal CT scan, assuming identical coverage, pitch, and tube rotation time? (1) Low kV-High mA (i.e., 80 kV, mAs 400) (2) High kV-Low mA (i.e., 140 kV, mAs 200)

Question 3 Match the answers with the questions. (There are more answers than questions.) (1) The effective dose of a routine head CT scan is

.

(2) The background radiation exposure of a North American adult is

.

(3) The American College of Radiology (ACR) recommends using a CT dose index volume (CTDIvol) of < for head imaging. (4) The ACR recommends using a CTDIvol of < Answers: (A) 25 mGy (B) 2 mSv (C) 5 mSv (D) 3 mSv (E) 75 mGy (F) 100 mGy

for abdominal imaging.

Test Questions

Question 4 This 18-year-old female presents with headaches. Four images are included from her head CT (Figures 9.1.1–9.1.4), along with a sagittal reconstruction (Figure 9.1.4). You suspect the appearance of the pituitary is due to: (1) a Rathke cyst. (2) a microadenoma. (3) a beam hardening artifact.

Figure 9.1.1

Figure 9.1.2

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Figure 9.1.3

Figure 9.1.4

Question 5 If you were to count them, there are over array.

individual detectors in a single row on a multidetector

(1) >10 (2) >50 (3) >500 (4) >1,000

Question 6 A chest CT has an effective dose that is (1) >10 (2) >20 (3) >50 (4) >100

times more than a PA and lateral chest X-ray.

Test Questions

Question 7 A routine head CT scan is performed with a CTDIvol ten times higher than a chest CT. As a result, the effective dose of a head CT is also much higher than that of a chest CT. True or False?

Question 8 The widespread use of modern multidetector scanners has substantially reduced patient dose from a single CT scan compared with that from single slice CT scanners. True or False?

Question 9 The mean energy, in keV, of the X-rays arising from an X-ray tube is: (1) the same value as peak kV (kVp). (2) 90% of the kvP. (3) 75% of the kVp. (4) 40% of the kVp. (5) 25% of the kVp.

Question 10 Automatic exposure control varies both the kV and mA of the X-ray tube to accommodate differences in thickness and attenuation of the body tissues. True or False?

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Answers 1. Correct answer: 2, 1, 3 Many will fi nd this surprising since we have been conditioned to believe that new means better in every way. When we consider CT critically, however, that is not necessarily true. Single-slice scanners have the highest geometric dose efficiency because the width of the X-ray beam is never wider than the detector. As a result, nearly all the X-rays that traverse the patient contribute to the image. Unlike single-slice CT scanning, however, the use of a multidetector array requires that the X-ray beam extend beyond the end detector rows so that every row sees the same number of X-rays as its neighbors. In order for the X-ray beam to cover all rows evenly, a portion of the beam must therefore extend beyond the detectors. This additional, and in one sense wasted, dose is described by overbeaming (see Chapter 2). The 16-slice scanner will be more dose efficient than a four-slice scanner, not because it has less overbeaming, but simply because requires fewer rotations to cover the same anatomy. By using helical technique with the 4 slice scanner there is not only the added dose from overbeaming but also the extra dose from “over-ranging”. In practice, the impact of overbeaming is considered insignificant on any CT scanners with more than 32 detector rows. 2. Correct answer: 1 The impact of a small increase or decrease of the kV used on the patient dose is much greater than a comparable change in mA. That is because the total dose increases as the square of kV but proportionally to mA. Whenever you choose to use a lower kV, as a result of the greater proportion of low-energy X-rays lost in peripheral tissue, an increase in mA will be necessary to provide comparable image quality. Even with that adjustment, however, the increase in dose from increasing the mA in thin patients and children will usually not be sufficient to offset the more substantial dose reduction from decreasing kV. If low-kV imaging is beneficial, you may be wondering why most sites use a kV of 120 instead of 80, 90, or even 100 kV for most head CT imaging. The choice of kV depends on the size and composition of the patient and tissues imaged. Even though it is preferred to use as low a kV as possible, it is still necessary to get enough X-rays to the detectors for adequate images. In fact, for very large patients and those with implanted metal, it may necessary to even increase the usual kV from 120 to 140 to provide adequate imaging. Low-dose but non-diagnostic imaging reminds me of a saying (if I may paraphrase a quote from Thomas Jefferson): The item you purchase on sale but don’t really want will prove to be the most expensive after all. Your goal should be to provide adequate imaging at the lowest possible dose but no one benefits from low-dose, non-diagnostic images. 3. Correct answers: 1-B, 2-D, 3-E, 4-A 4. Correct answer: 3 The sagittal reconstruction (Figure 9.1.4) shows a low-attenuation area within the pituitary that corresponds to the low attenuation evident on the axial image within the sella (Figure 9.1.3). This is due entirely to beam hardening caused by the bony margins of the sella. In this case, the pituitary was imaged in the direct coronal view, and that image shows a prominent but normal pituitary gland for a woman of her age.

Test Questions

5. Correct answer: 3 In a single row, there are usually 700–1,000 individual detectors. Multiply that by the 128 detector rows or more that are routine on many new CT scanners, and you not only have a lot of detectors but also a stunning amount of data streaming to the computer on each and every rotation. 6. Correct answer: 3 A chest X-ray has an effective dose of 0.1 mSv, whereas a chest CT has an effective dose of 7–10 mSv. 7. Correct answer: False Since the effective dose reflects both the magnitude of the absorbed dose as well as the sensitivity of the involved tissues to radiation, the chest CT has a 3–5 times higher effective dose than does a head CT. That is because the stochastic risks of radiation to the brain are considered to be much lower than the same or even less dose applied to breast tissue and esophagus. 8. Correct answer: False It is somewhat counterintuitive, but the late-generation single-slice axial scanners were quite efficient in their use of radiation. Wider detector arrays with helical imaging and narrow detector collimation have improved imaging speed and make many new exams possible but do not necessarily reduce dose for comparable scans. The use of automatic exposure control (AEC) in many cases, along with careful attention to the scan technique, will significantly mitigate any of the dose inefficiencies inherent with multidetector CT imaging. There is great hope for newer reconstruction techniques to bring patient dose down even lower. 9. Correct answer: 4 The energy of the X-rays emitted by the anode is expressed in keV, thousand electron Volts, not as kVp. Yet the two values are linked. These X-rays have a range of values. Although none of these will have a higher energy than that predicted by the kVp, depending on filtration the mean of the all the X-ray energies will be about 30–50% of that of the highest. 10. Correct answer: False Automatic exposure control on most scanners modulates only the mA. Although kV undoubtedly has a significant impact on dose, the scanner software will not vary it in response to the tissues examined. For those of you familiar with camera operation, think of AEC like the automatic aperture mode on a camera. Although both shutter speed and aperture can be used together and independently to influence the exposure of a photograph, in aperture mode only the size of the aperture varies and shutter speed always stays the same. Some modern CT scanners offer a method to alter both kV and mA but that requires the use of a ‘‘contrast index’’ since noise index or equivalent mA reference values will need to be altered as kV changes.

THE END

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INDEX Index entries followed by t indicate a table; by f indicate a figure. absorbed dose, 37 absorption efficiency, 23 accentuation artifact, 96 acute stroke, 179 acute subdural hemorrhage, 185, 185f AEC. See automatic exposure control afterglow, 23 air in arterial system, 149, 150f cranioplasty fl ap, trapping, 168f–169f fat compared to, 171, 172f peristalsis and, 213, 213f in subarachnoid space, 148–51, 148f–151f in veins, 149, 151f ALARA. See As Low As Reasonably Achievable algebraic reconstructions, 18, 18f amperage principles, 46. See also milli-ampere angiography unit fl at plate detector, 22f angular adjustment, 50 anisotropic voxels, 16f arachnoid granulation, 178–80, 178f–180f arms raised, photon starvation avoided with, 203, 204f artifacts accentuation, 96 beam hardening, 133–35, 133f, 135f–136f, 141, 143f blooming, 67–68, 68f, 97, 98f blurring, 109–11, 177 chest wall motion artifact, 80–81 craniectomy, 130–31, 130f–132f CTA, 67–68, 68f cupping, 135 from densely calcified coronary vessels, 90–92, 90f–92f dental amalgam, 140–41, 140f–144f gas bubble motion, 212–13, 212f–213f Hounsfield, 143f incomplete projection, 205–6, 205f–206f metallic, 68–69, 95–98, 95f–98f, 198–99, 198f posterior fossa, 140, 140f from prospective triggering, 93–94, 93f, 94f from random patient motion, 75–78, 75f–78f, 126–27, 126f–129f respiratory motion differentiated from cardiac motion, 79–82, 79f–82f in scanograms, 126–27, 126f–129f stair-step, 109–11

streak, 134 susceptibility, 134 volume averaging, 114–18, 114f–118f, 152–53, 152f–153f, 223–24, 223f–224f windmill, 130–31, 130f–132f artifacts, body CT barium, 200, 200f detector array, 207–8, 207f–208f gas bubble motion artifact, 212–13, 212f–213f incomplete projection, 205–6, 205f–206f metallic artifacts, 198–99, 198f photon starvation, 202–3, 202f–204f random patient motion, 209–10, 209f–211f volume averaging artifacts, 223–24, 223f–224f artifacts, cardiac imaging densely calcified coronary vessels, 90–92, 90f–92f ECG editing, 99–101, 99f–100f heart rate increases, 102–4, 102f–104f irregular heart rates, 109–11, 109f–111f metal hardware, 95–98, 95f–98f prospective triggering, 93–94, 93f, 94f PVCs, 87–89, 87f, 88f random patient motion, 75–78, 75f–78f respiratory motion differentiated from cardiac motion, 79–82, 79f–82f suboptimal image contrast, 105–8, 105f–106f, 108f temporal resolution, 72–74, 72f–73f tube current modulation and cardiac CTA, 83–86, 83f–86f artifacts, neuro CT beam hardening, 133–35, 133f, 135f–136f, 141, 143f dental amalgam, 140–41, 140f–144f detector calibration error, 130–31, 130f–132f from gantry angulation, 141 in-plane skull fractures, 122, 122f–125f from MRI location compared to CT, 145–46, 145f photon starvation, 137–39, 137f–139f random patient motion causing, 126–27, 126f–129f subarachnoid hemorrhage attenuation values, 119–21, 119f–121f volume averaging, 114–18, 114f–118f, 152–53, 152f–153f As Low As Reasonably Achievable (ALARA), 39 atrial fibrillation, 100, 109 attenuation values, 12–13, 15f arachnoid granulation, 178–80, 178f–180f CSF in empty sella, 134

234

INDEX attenuation values (Cont.) metallic artifacts and, 95–98, 95f–98f pixel size, 115 SMV, 218–19, 218f–220f subarachnoid hemorrhage and, 119–21, 119f–121f of subdural hemorrhage, 165–66 voxel, 115 automatic exposure control (AEC), 32, 203, 231 for chest CT, 62–63, 64f important reminders for, 50–51 mA modulation with, 229, 231 radiation dose reduction and, 50, 52 shields and, 51–52 tube current modulation adjustments for radiation dose, 61–62 axial imaging, 26f Chiari I malformation problems, 155, 156f detector calibration error artifacts, 131, 132f helical imaging compared to, 6–10, 42 limits of, 182 misregistration in, 7, 9f motion effects on, 127, 129f for MRI, 145 prospective triggering and, 60–61 random patient motion and, 76–77, 77f rotation time in, 29 scan time for, 7 back-projection, 8f, 52. See also fi ltered back-projection fi lters for, 18 history of, 18, 18f, 19f barium artifacts, 200, 200f beam collimation, 10–12, 28 efficiency of, 42, 43f beam hardening, 46 artifacts, 133–35, 133f, 135f–136f, 141, 143f detector array sensitivity to, 208 infarct obscured by, 183, 183f photo starvation compared to, 138 principles of, 133–34 pseudo-enhancement from, 217 beta-blockers, 68 bismuth, 51 blood, in extra-axial space, 166 blood hematocrit levels, 179 blooming artifact in CTA, 67–68, 68f high-density structures and, 96, 96f blurring artifacts, 109–11, 177 body CT artifacts barium, 200, 200f detector arrays unevenly calibrated or malfunctioning causing, 207–8, 207f–208f gas bubble motion artifact, 212–13, 212f–213f incomplete projection, 205–6, 205f–206f metallic artifacts, 198–99, 198f photon starvation, 202–3, 202f–204f random patient motion, 209–10, 209f–211f

body CT pitfalls false-positive venous thrombosis diagnosis, 222, 221f–222f renal cyst pseudo-enhancement, 216–17, 216f–217f SMV normal low attenuation values, 218–19, 218f–220f volume averaging artifacts, 223–24, 223f–224f bolus tracking, 107 brain atrophy, 182 brain scans, gantry angulation issues with, 26, 27f brain tumor, 189–90, 189f–191f infarct misdiagnosis of, 189–90, 189f–191f metastatic, 190, 191f breath-holding procedure, 59, 81–82, 107, 209 bullets and metal artifact, 97f calcifications. See densely calcified coronary vessels cancellation of diagnostic study, 220 cardiac cycle phases, 103–4 cardiac imaging of coronary vessels, 66–67, 67f CTA studies with, 67–68, 67f densely calcified coronary vessel artifacts in, 90–92, 90f–92f dual-source CT scanning improvements for, 59–60, 60f ECG editing artifacts in, 99–101, 99f–100f ECG synchronization for, 60–61 heart rate increases artifacts in, 102–4, 102f–104f irregular heart rate artifacts in, 109–11, 109f–111f metal hardware artifacts in, 95–98, 95f–98f nonsynchronization issues in, 66–69 normal rhythm in, 56f partial scan for, 59, 59f photon starvation in, 69 prospective triggering artifacts in, 93–94, 93f, 94f PVC artifacts in, 87–89, 87f, 88f radiation dose optimization for, 61–66, 62f random patient motion artifacts in, 75–78, 75f–78f respiratory motion differentiated from cardiac motion artifacts in, 79–82, 79f–82f segmented acquisition for, 58–59 SNR needs in, 62 suboptimal image contrast artifacts in, 105–8, 105f–106f, 108f temporal resolution and, 56–57, 56f temporal resolution artifacts in, 72–74, 72f–73f tube current modulation and cardiac CTA artifacts in, 83–86, 83f–86f tube current modulation matched to cycle of, 64 cardiac motion ECG suppressing, 76 respiratory motion, 79–82, 79f–82f cavernoma, 157–58, 157f–160f cavernous sinus, 172f central nervous system disease, MRI over CT for, 193, 193f–194f cerebellar infarct, 181–83, 181f–183f cerebellopontine angle cistern, 194f cerebellum, 120 hemorrhage of, 121f

Index cerebrospinal fluid (CSF) attenuation values in empty sella of, 134 in subdural space, 185, 187f chest CT AEC for, 62–63, 64f chest X-ray effective dose compared to, 228, 231 effective dose in head CT compared to, 228, 231 respiratory motion differentiated from cardiac motion in, 79–82, 79f–82f chest wall motion artifact, 81 Chiari I malformation, 155, 156f children. See pediatric CT imaging chronic subdural hemorrhage, 185, 185f chronic subdural, wide subarachnoid space mistaken for, 185, 187f cobalt alloy clip, 139f coiled aneurysms, photon starvation and, 138–39, 144f collimators, 42 comb (plastic) fragments, 170–71, 170f–173f complicated cyst, 216 computed axial tomography (CAT and CT scanning), 4 computed tomography (CT), 2. See also dual-energy CT scanning; dual-source CT scanning; multidetector CT algebraic reconstructions for, 18, 18f cone beam imaging, 20–24, 22f contrast-enhanced, 174–77, 174f–177f dose measures for, 36–37 EMI development of, 4–5 Hounsfield’s advances with, 4–5, 7f indications for, 41–42 MRI compared to, 192, 193, 194f, 195f MRI for central nervous system disease over, 192, 193f–194f neuro CT artifacts comparing MRI to, 145–46, 145f pediatric, 52–53 resolution and voxel size in, 13 speed, 177 spiral, 7–8 third-generation units, 9f, 23f time delay, 29 computed tomography angiography (CTA), 29 blooming artifact in, 67–68, 68f cardiac imaging studies with, 67–68, 68f densely calcified coronary vessels and, 90–92, 90f–92f future of, 30 retrospective gating and, 101 SNR modification in, 73–74 speed of, 29–30 temporal resolution with dual-source CT scan for large patients in, 72, 73 tube current modulation and cardiac, 83–86, 83f–86f computed tomography dose index (CTDI), 36–37, 226 computed tomography dose index volume (CTDIvol), 36 computed tomography venogram (CTV), 179 cone beam, 19 cone beam imaging, 20–24, 22f contrast. See image contrast contrast-enhanced CT scan, CTA compared to, 174–77, 174f–177

contrast index, 231 contrast opacification, 106 contrast staining, 162, 164f Cormack, Allan, 3, 5 coronal reconstruction for arachnoid granulation, 178f, 179 for posterior fossa, 182, 182f for transverse sinus asymmetry, 179, 180f window and level settings for, 166, 170f coronary artery bypass graft (CABG), 96 coronary vessels, 66–67, 67f densely calcified, 90–92, 90f–92f cortical ribbon, 185, 186f craniectomy, 130–31, 130f–132f cranioplasty, air trapped in, 168f–169f CT. See computed tomography CTA. See computed tomography angiography CTDI. See computed tomography dose index CTDIvol. See computed tomography dose index volume CTV. See computed tomography venogram cupping artifact, 135 Curie, Marie, 39 Cyst and CT complicated, 216 pseudo enhancement of renal, 216–17, 216f–217f data gaps, 110–11, 111f densely calcified coronary vessels, 90–92, 90f–92f dental amalgam artifacts, 140–41, 140f–144f detector arrays beam hardening, 208 in multidetector CT, 228, 231 reconstructions and, 15–17, 18–20 septa in, 21 solid-state, 207 unevenly calibrated or malfunctioning, 207–8, 207f–208f xenon gas, 207 detector calibration error artifacts, 130–31, 130f–132f detector collimation, 16 narrow, 131 slice thickness and, 49 thin, 49 deterministic effects, of radiation dose, 37, 38f developmental venous anomaly, 160f diastolic phase end, 93–94 heart rate increases shortening, 103 DLP. See dose length product dose. See radiation dose dose efficiency, 23 of various scanners, 226, 230 dose length product (DLP), 39–40 dual-energy CT scanning, 49f benefits of, 29, 31f future of, 30 virtual noncontrast image processing and, 162, 163f dual-source CT scanning gantry rotation time for, 60

235

236

INDEX dual-source CT scanning (Cont.) prospectively triggered high-pitch helical acquisition for, 66, 67f scan acquisition speed in, 29, 30f temporal resolution for cardiac imaging improvements with, 59–60, 60f temporal resolution for CTA in large patients with, 72–75, 72f–73f ECG. See electrocardiogram ectopic beats, 86 effective dose, 226 in chest CT compared to head CT, 228, 231 of chest CT to chest X-rays, 228, 231 DLP conversion to, 39–40 measuring, 38 Monte Carlo simulation and, 39 effective mAs (mAseff), 46–47 Electrical and Musical Industries (EMI), 4–5, 6, 10 electrocardiogram (ECG) artifacts problems with editing of, 99–101, 99f–100f cardiac motion suppressed with, 76 data gaps from PVC editing disruption in, 110–11, 111f lead contact and, 100–101 PVC editing in, 87–89, 87f, 88f retrospective gating and editing of, 88, 88f, 101 synchronization, 60–61, 99–101, 99f–100f tube current modulation and, 62–93, 64f Elscint, 11 EMI. See Electrical and Musical Industries end diastolic imaging, 58 end diastolic phase, 93–94 end systole phase, 93–94 extra-axial space blood in, 166 wide subarachnoid space mistaken for chronic subdural space in, 185, 187f eye and dose, gantry angulation, 26–27

isotropic voxels and, 26, 26f neuro CT artifacts from, 140 radiation dose to eye in, 26–28 gantry rotation time for dual-source CT scanning, 59–60 limitations to, 56, 57f partial scan and, 59, 59f segmented acquisition and, 58–59 temporal resolution and, 56–57, 57t temporal resolution improvements for, 57–60, 58f gas bubble motion artifact, 212–13, 212f–213f gastric lumen, 212, 212f Gelfoam, 171, 172f geometric efficiency, 23 glucagon, 213 gold, 199 Gray (Gy), 37 Gray, Louis Harold, 37 gunshot injury to brain, 138–39, 138f–139f, 141f Gy. See Gray

fat, air compared to, 171, 172f fat emboli, 148–49, 148f FBP. See fi ltered back-projection fi ltered back-projection (FBP), 8f iterative reconstruction, compared to, 66, 66f metallic artifacts and, 96, 97f for reconstruction, 65–66 fi lters for back-projection, 18 dark, 177 types of, 18 windowing compared to, 17f fl at-panel detector, 20, 23. See also cone beam imaging flow artifacts, venous thrombosis compared to, 222, 222f four-dimensional CT scanning, for respiratory motion, 82, 82f Frank, Gabriel, 18, 19f

hair loss, 37, 38f head CT scan, 36 effective dose in chest CT compared to, 2289, 231 heart rates cardiac imaging artifacts with irregular, 109–11, 109f–111f diastolic phase shortening by increase in, 103 motion-free imaging and artifact issues with increases in, 102–4, 102f–104f prospective triggering and variation of, 94, 94f radiation dose reduction and variability of, 110 segmented acquisition and temporal resolution for elevated, 104, 104f tube current modulation and irregular, 84 helical imaging axial imaging compared to, 6–10, 42 for detector calibration error artifacts, 130–31, 130f–132f interpolation and, 10 interpolation for reconstruction in, 20, 21f prospectively triggered high-pitch helical acquisition for, 66, 67f retrospective gating and, 60–61 rotation time in, 29 slip rings for, 8, 11f X-ray beam trajectory in, 21f hematocrit levels, blood, 179 hemorrhage, 115, 118f attenuation values and subarachnoid, 119–21, 119f–121f cerebellar, 121f missed subdural, 165–66, 165f–169f, 184–85, 184f–188f high-pitch scanning, 76, 81 Hounsfield, Godfrey, 3–4 CT advances of, 4–5, 7f Hounsfield artifact, 143f Hounsfield unit (HU), 4

gamma rays, 4 gantry angulation image display and, 25–27

image contrast basis of, 12–13 kV increases and, 106

Index minimizing suboptimal, 105–7, 105f–106f, 108f time delay and CT, 175, 175f window and level settings for, 166, 167f–169f image display, gantry angulation and, 25–27 “Image Gently,” 53 implanted metal, 198 incomplete projection artifact, 205–6, 205f–206f infarct beam hardening obscuring, 183, 183f brain tumor misdiagnosed as, 189–90, 189f–191f cerebellar, 181–83, 181f–183f wide sulcus distinguished from, 181–83, 181f–183f inferior vena cava (IVC), 107, 108f in-plane location of skull fractures, 122, 122f–125f, 124 interpolation, 9 helical imaging and, 10 for helical imaging reconstruction, 19, 21f iodine, photoelectric effect with, 29, 31f IR. See iterative reconstruction irregular heart rates. See heart rates isotropic voxels gantry angulation and, 26, 26f reconstructions with, 15–17, 16f iterative reconstruction (IR), 24, 25f FBP compared to, 66, 66f for metallic artifacts, 97 radiation dose reduction and, 52 Kalender, Willi, 8 kernels, 18. See also fi lters fi lter terminology choice compared, 50 kilovolt (kV) choosing strength of, 226, 230 decreasing, 47–48, 49f image contrast increasing, 106 increasing, 47, 48f principles of, 45–46 radiation dose relationship to, 47 SNR with modification of, 73–74 lens dose, 146 level settings, image contrast and, 166, 167f–169f Litvinenko, Alexander, 37 longitudinal axis collimation, 81 lumbar myelogram, 161–64, 161f–164f mA. See milli-ampere magnetic resonance imaging (MRI), 13, 41 axial imaging for, 145 of cavernoma, 158, 158f–160f for central nervous system disease over CT, 192–93, 192f–195f CT compared to, 192–93, 194f, 195f metallic artifact risks for, 198 neuro CT artifacts comparing CT to, 145–46, 145f magnetic resonance venogram (MRV), 179 mAs. See milli-ampere seconds mAseff. See effective mAs matrix size, pixel size relationship with, 115

meningioma, 177, 175f metallic artifact reduction (MAR), 96–97 metallic artifacts, 68–69. See also photon starvation attenuation values and, 95–98, 95f–98f bullets, 97f, 137–39, 137f–139f, 141f dental amalgam artifacts and, 140–41, 140f–144f FBP and, 96, 97f IR for, 97 mGy. See milli-Gray middle cerebral artery (MCA), 189 milli-ampere (mA) AEC modulating, 229, 231 kV changes compared to, 226, 230 principles of, 45–46 radiation dose linear relationship to, 46 rotation time and, 45 SNR with modification of, 73–74 milli-ampere seconds (mAs), 45 milli-Gray (mGy), 37 milli-Sieverts (mSv), 38 misregistration, in axial imaging, 7–8, 9f Monte Carlo simulation, 39 motion. See also random patient motion axial imaging and, 127, 128f cardiac, 76, 79–82, 79f–82f gas bubble motion artifact, 212–13, 212f–213f overscanning for, 76 random patient, 75–78, 75f–78f, 126–27, 126f–129f, 192–93, 192f–195f respiratory, 79–82, 79f–82f mSv. See milli-Sieverts multidetector CT detector arrays in, 228, 231 fl at-panel detector compared to, 21, 23 history and advancements with, 10–11 overbeaming in, 42, 43f slice thickness of, 15 320-row, 11, 13f widespread use of, 229, 231 myelographic contrast, 162, 162f neck imaging, time delay for soft-tissue, 175, 177f neuro CT artifacts beam hardening, 133–35, 133f–136f, 141, 143f dental amalgam, 140–41, 140f–144f detector calibration error, 130–31, 130f–132f from gantry angulation, 141 in-plane location of skull fractures in, 122, 122f–125f, 124 from MRI compared to CT, 145–46, 145f photon starvation, 137–39, 137f–139f random patient motion causing, 126–27, 126f–129f subarachnoid hemorrhage attenuation value, 119–20, 119f–121f volume averaging, 114–15, 114f, 116–18f, 152–53, 152f–153f neuro CT pitfalls air in subarachnoid space, 148–49, 148f–151f air misconceptions, 170–72, 170f–173f

237

238

INDEX neuro CT pitfalls (Cont.) arachnoid granulation, 178–79, 178f–180f brain tumor misdiagnosed as infarct, 189–90, 189f–191f cavernoma, 157–58, 157f–160f CTA disadvantages compared to contrast enhanced CT scan, 174–75, 174f–177f, 177 infarct distinguished from wide sulcus, 181–83, 181f–183f lumbar myelogram, 161–64, 161f–164f missed subdural hemorrhage, 165–66, 165f, 167f–169f random patient motion, 192–93, 192f–195f ”rule out stroke,” 189–90, 189f–191f sellar masses, 154–55, 154f–156f volume averaging artifact, 152–53, 152f–153f “noise index,” 32 occipital lobes, 135, 136f optics, 26 optimal dose, 40, 41f orbital roof, 115 osteomyelitis, 168f overbeaming, 26 in multidetector CT, 42, 43f overranging, 42–44, 44f overscanning, 76, 210 PAC. See premature atrial contraction Pacchionian granulations, 179 partial scan, 59, 59f patient comfort, 210 patient motion. See random patient motion pediatric CT imaging radiation dose and, 52–53 subarachnoid space enlargement in, 185, 188f penetrating injury, wood and, 171, 173f penumbra, 42 peristalsis, 213, 213f petrous bone, 115 photoelectric effect, 29, 31f photon starvation arms raised for avoiding, 202–3, 202f–204f beam hardening compared to, 138 body CT artifacts with, 202–3, 202f–204f in cardiac imaging, 68–69 coiled aneurysms and, 138, 144f neuro CT artifacts from, 137–39, 137f–139f titanium aneurysm clips and, 138, 139f pitch choosing, 44, 45f, 45f high-pitch scanning, 76, 81 prospectively triggered high-pitch helical acquisition and, 66, 67f scan acquisition speed and, 28, 28f pitfalls, body CT false-positive venous thrombosis diagnosis, 221–22, 221f–222f renal cyst pseudo-enhancement, 216–17, 216f SMV normal low attenuation values, 218–19, 218f–220f pitfalls, neuro CT

air in subarachnoid space, 148–49, 148f–151f air misconceptions, 170–72, 170f–173f arachnoid granulation, 178–79, 178f–180f brain tumor misdiagnosed as infarct, 189–90, 189f–191f cavernoma, 157–58, 157f–160f CTA disadvantages compared to contrast enhanced CT scan, 174–75, 174f–177f, 177 infarct distinguished from wide sulcus, 181–83, 181f–183f lumbar myelogram, 161–64, 161f–164f missed subdural hemorrhage, 165–66, 165f, 167f–169f, 184–85, 184f–188f random patient motion, 192–93, 192f–195f ”rule out stroke,” 189–90, 189f–191f sellar masses, 154–55, 154f–156f volume averaging artifact, 152–53, 152f–153f pituitary gland, beam hardening and, 227, 227f–228f, 230 pituitary macroadenoma, 155f pixel size, 13 attenuation values assigned to, 115 matrix size relationship with, 115 plastic, 171, 171f pneumocephalus, 151f pons, 135, 135f encephalomalacia in, 143f posterior fossa, 120 artifacts, 140, 140f coronal reconstruction helpful for imaging, 182, 182f CT problems imaging, 182 premature atrial contraction (PAC), 94 premature ventricular contraction (PVC) data gaps in disabled ECG synchronization from, 109–10, 111f ECG editing for, 87–89, 87f, 88f prospectively triggered high-pitch helical acquisition, 66, 67f prospective triggering artifacts from, 93–94, 93f, 94f axial imaging and, 60–61 heart rate variation and, 94, 94f radiation dose reduction with tube current modulation and, 84–86, 84f proximal left middle cerebral artery, 164f pseudo-enhancement, renal cyst, 216–17, 216f PVC. See premature ventricular contraction QRS complex, 103 radiation cost, 40 radiation dose AEC and reduction of, 50–51 cardiac imaging optimizing, 61–66, 62f deterministic effects of, 37, 38f to eye in gantry angulation, 26–27 heart rate variability and reduction of, 110 IR and reduction of, 52 kV relationship to, 47 mA linear relationship to, 46 measures for, 36–37

Index optimal, 40, 41f overranging and, 42–44, 45f pediatric CT imaging and, 52–53 reduction techniques for, 40–41 stochastic effects of, 38–39 terminology for, 36 tube current modulation adjustments without AEC for, 63–64 tube current modulation and prospective triggering reducing, 84–86, 84f tube current modulation linear changes with, 62 tube voltage square changes with, 62 radiation sickness, 37 radioactivity, 39 random patient motion artifacts from, 75–78, 75f–78f, 126–27, 126f–129f axial imaging and, 76, 78f body CT artifacts from, 209–10, 209f–211f neuro CT artifacts from, 126–27, 126f–129f overscanning for, 76 preventing, 76, 78f skull fractures mistaken for, 126–27, 126f–129f tracking, 76–77, 78f reconstruction. See also coronal reconstruction algebraic, 18, 18f beam hardening artifacts appearance in, 141, 143f detector arrays and, 18–19 FBP for, 65–66 interpolation for helical imaging, 19, 21f isotropic voxels for, 15–17, 16f iterative, 24, 25f kymogram-correlated, 100 multiphase, 88–89 skull fractures and sagittal, 124, 124f–125f windowing for, 17f, 18 recovery time, 23 “reference mA,” 32 renal cyst pseudo-enhancement, 216–17, 216f rescheduling examination, 200–201 resolution, 13 respiratory motion cardiac motion differentiated from, 79–82, 79f–82f four-dimensional CT scanning for, 82, 82f retrospective gating, 60–61 CTA and, 101 ECG editing in, 88–89, 88f, 101 right coronary artery (RCA), 110 right-sided subdural hemorrhage, 183 rings. See slip rings Roentgen, Wilhelm, 2 rotation time, 23, 28. See also gantry rotation time in axial imaging, 29 in helical imaging, 29 mA and, 45 “rule out infarct,” 145 “rule out stroke,” 189–90, 189f–191f sagittal reconstruction, of skull fractures, 124, 125f

scan acquisition speed in dual-source CT scanning, 29, 30f factors determining, 28 pitch and, 28–29, 28f scanograms, 50, 51f artifacts in, 126–27, 126f–129f “scan thin, view thick” principle, 16 segmented acquisition, temporal resolution and, 58–59, 104, 104f septa, 21 shields, 51–52 Sievert, Rolf, 38 Sieverts, 38 signal-to-noise ratio (SNR), 16, 48–49 cardiac imaging needs with, 66 CTA modification of, 73–74 kV and mA modification for, 74 optimizing, 61–66 tube current modulation increasing, 85–86, 86f skull base masses, 154–55, 154f–156f skull fractures in-plane location of, 122, 122f–125f random patient motion mistaken for, 126–27, 126f–129f sagittal reconstruction of, 124, 125f slice thickness and, 116f–118f slice thickness detector collimation and, 48–49 historical improvements in, 13–15 of multidetector CT, 15 skull fractures and, 116f–118f SNR and, 16 voxel clarity and, 115, 116f slip rings artifacts from, 130–31, 130f–132f, 207–8, 208f for helical imaging, 8, 12f for television antennas, 11f SMV. See superior mesenteric vein SNR. See signal-to-noise ratio soft-tissue neck imaging, time delay for, 175, 177f solid-state detectors, 207 spectral distribution, 61 speed problems, for CT, 177 spiral CT, 7–8. See also helical imaging splenic infarct, 204f stair-step artifacts, 110 step-and-shoot. See axial imaging stochastic effects, of radiation, 38–39 streak artifact, 134 stroke, rule out, 189–90, 189f–191f subarachnoid hemorrhage attenuation values representing, 119–20, 119f–121f cerebellar hemorrhage mistaken for, 121f hydrocephalus and, 162, 162f–163f tentorial, 179, 180f subarachnoid space air in, 148–49, 148f–151f pediatric CT imaging showing enlargement of, 185, 188f virtual noncontrast image processing for, 162, 163f–164f

239

240

INDEX subcutaneous fat, 171, 172f subdural hemorrhage acute, 185, 185f attenuation values of, 166 chronic, 185, 185f cortical ribbon in, 185, 186f missed, 165–66, 165f, 167f–169f, 184–85, 184f–188f right-sided, 183 superior mesenteric vein (SMV), 218–19, 218f–220f super vena cava (SVC), 107, 108f SVC. See super vena cava sylvian fi ssure, 160f, 190 systole phase, end, 94 table speed, 28 television antennas, 11f temporal resolution cardiac imaging and, 56–57, 56f for CTA in large patients with dual-source CT scan, 72–74, 72f–74f dual-source CT scanning improvements for, 59–60, 60f gantry rotation time and, 56–57, 57t gantry rotation time improvements for, 57–60, 58f partial scan and, 59, 59f segmented acquisition and, 58–59, 104, 104f test answers, 230–31 test questions, 225–29 thrombus, 179 time delay, 29 with contrast-enhanced CT scans, 175, 176f image contrast increasing, 177, 175f for soft-tissue neck imaging, 175, 177f titanium, 199 titanium aneurysm clips, 138, 139f tomography. See also computed tomography mechanics of, 5f X-ray, 3 “translate” movement, 7, 8f transverse sinus, 179, 179f asymmetry, 179, 180f tube current modulation, 50. See also automatic exposure control cardiac CTA and, 83–86, 83f–86f cardiac imaging cycle matched to, 65 ECG and, 64–65, 65f irregular heart rates and, 84 photon generation quantity determined by, 61–62, 62f radiation dose linear changes with, 62 radiation dose reduction with prospective triggering and, 84–86, 84f radiation dose without AEC adjusting, 63–64 SNR increases with, 85–86, 86f tube voltage radiation dose changes with square of, 62 spectral distribution quality determined by, 61, 62f

umbilical ring and chain, 198–99, 198f venous air, 149, 151f venous thrombosis, 179 body CT pitfalls of false-positive diagnosis of, 221–22, 221f–222f flow artifacts compared to, 222, 222f ventricular diastole, 102–4 ventricular tachycardia, 100 virtual noncontrast image processing, 162, 163f–164f Volta, Alessandro, 46 voltage. See also kilovolt measuring, 46 principles of, 45–46 X-ray beam energy and, 74 volume averaging artifacts body CT pitfalls with, 223–24, 223f–224f neuro CT pitfalls with, 114–15, 114f, 116–18f, 152–53, 152f–153f voxels, 115 anisotropic, 16f attenuation values, 115 isotropic, 15–17, 16f, 26, 26f resolution and size of, 13 slice thickness and clarity of, 115, 116f wide subarachnoid space, chronic subdural space mistaken for, 185, 187f wide sulcus, infarct distinguished from, 181–83, 181f–183f windmill artifact, 130–31, 130f–132f windowing fi lters compared to, 17f reconstruction using, 17f, 18 for wood in penetrating injury, 171, 173f window settings, image contrast and, 166, 167f–169f wood appearance variations of, 171 windowing for penetrating injury with, 171, 173f xenon gas detectors, 207 X-rays. See also beam hardening chest CT effective dose compared to chest, 229, 231 depth limitations of, 4f discovery and early use of, 2, 2f energy composition of, 134 helical, 10f helical imaging trajectory of, 21f mean energy from anode of, 229, 231 photoelectric effect with iodine and, 29, 31f scattering problem of, 2–3, 21 tissue contrast limitations of, 2, 3f tomography, 3 voltage and beam energy of, 74