ACCF/AHA/HRS/SCAI FLUOROSCOPY CLINICAL COMPETENCE STATEMENT
ACCF/AHA/HRS/SCAI clinical competence statement on physician knowledge to optimize patient safety and image quality in fluoroscopically guided invasive cardiovascular procedures
A report of the American College of Cardiology Foundation/American Heart Association/American College of Physicians Task Force on Clinical Competence and Training
John W. Hirshfeld, Jr, MD, FACC, FAHA, FSCAI, Chair, Writing Committee Member,
Stephen Balter, PhD, FACR, FAAPM, FSIR, Writing Committee Member,
Jeffrey A. Brinker, MD, FACC, FSCAI, Writing Committee Member,
Morton J. Kern, MD, FACC, FAHA, FSCAI, Writing Committee Member*,
Lloyd W. Klein, MD, FACC, FAHA, FSCAI, Writing Committee Member ,
Bruce D. Lindsay, MD, FACC, FAHA, Writing Committee Member ,
Carl L. Tommaso, MD, FACC, FAHA, FSCAI, Writing Committee Member,
Cynthia M. Tracy, MD, FACC, FAHA, Writing Committee Member*,,
Louis K. Wagner, PhD, FACR, FAAPM, Writing Committee Member,
Mark A. Creager, MD, FACC, FAHA, Chair, Task Force Member,
Michael Elnicki, MD, FACP, Task Force Member,
John W. Hirshfeld, Jr, MD, FACC, FAHA, Task Force Member,
Beverly H. Lorell, MD, FACC, FAHA, Task Force Member,
George P. Rodgers, MD, FACC, Task Force Member,
Cynthia M. Tracy, MD, FACC, FAHA, Task Force Member and
Howard H. Weitz, MD, FACC, FACP, Task Force Member
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Table of contents
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Preamble......2260
- I Introduction and Background......2260
- II The Physics and Nature of X-Radiation......2262
The Nature of X-Radiation......2262
X-Ray Generation......2262
X-Ray Spectra......2262
Mechanisms of X-Ray Absorption......2262
X-Ray Dose......2263
Measurements of Radiation......2263
- III Principles of X-Ray Image Formation......2263
X-Ray Image Generation......2263
Parameters That Affect X-Ray Image Formation......2263
- IV The Operation of an X-Ray Cinefluorographic Unit......2266
Overview......2266
X-Ray Generation......2266
X-Ray Exposure Modulation......2268
Video Image Capture......2268
Image Display and Processing......2269
Optimizing the Exposure Parameters That Determine Image Quality......2270
Fluoroscopy Dose Management Issues......2270
Acquisition Dose Management Issues......2271
Digital Image Subtraction......2271
- V Determinants of Patient X-Ray Dose......2271
Measurements of Patient Dose......2271
Dose (Air Kerma) at the IRP......2271
DAP......2271
Value of Dose Monitoring......2272
Dose at the IRP Monitoring......2272
DAP Benchmarking Data......2272
Factors That Influence Patient Absorbed Dose......2273
Equipment-Related Factors......2273
Patient-Related Factors......2273
Procedure Conduct Factors......2273
Beam Collimation......2274
Input Dose and Frame Rate......2274
- VI Patient Effects of X-Ray Exposure......2274
Deterministic Effects......2275
Cataract......2276
Stochastic Effects......2276
- VII Radiation Risks From Typical Invasive Cardiovascular Procedures......2277
Clinical Risk-Benefit Ratio......2277
General Considerations for Deterministic Risk......2277
General Considerations for Stochastic Risk......2278
Multiple Procedure Considerations......2278
- VIII Physician Responsibilities to Patients......2278
Patient Education and Consent......2278
Procedural Dose Management......2278
- IX Recommended Radiation Safety Curriculum for Physicians Who Perform Invasive Cardiac Procedures......2279
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Preamble
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The granting of clinical staff privileges to physicians is a primary mechanism used by institutions to uphold the quality of care. The Joint Commission on Accreditation of Healthcare Organization (JCAHO) requires that the granting of continuing medical staff privileges be based on assessments of applicants against professional criteria specified in the medical staff bylaws. Physicians themselves are thus charged with identifying the criteria that constitute professional competence and with evaluating their peers accordingly. Yet the process of evaluating physicians' knowledge and competence is often constrained by the evaluator's own knowledge and ability to elicit the appropriate information, problems compounded by the growing number of highly specialized procedures for which privileges are requested.
The American College of Cardiology/American Heart Association/American College of Physicians (ACC/AHA/ACP) Task Force on Clinical Competence was formed in 1998 to develop recommendations for attaining and maintaining the cognitive and technical skills necessary for the competent performance of a specific cardiovascular service, procedure, or technology. These documents are evidence-based, and where evidence is not available, expert opinion is utilized to formulate recommendations. Indications and contraindications for specific services or procedures are not included in the scope of these documents. Recommendations are intended to assist those who must judge the competence of cardiovascular health care providers entering practice for the first time and/or those who are in practice and undergo periodic review of their practice expertise. The assessment of competence is complex and multidimensional; therefore, isolated recommendations contained herein may not necessarily be sufficient or appropriate for judging overall competence.
The ACC/AHA/ACP Task Force on Clinical Competence makes every effort to avoid any actual or potential conflicts of interest that might arise as a result of an outside relationship or a personal interest of a member of the writing panel. Specifically, all members of the writing panel were asked to provide disclosure statements of all such relationships that might be perceived as real or potential conflicts of interest. These statements were reviewed by the ACC/AHA/ACP Task Force on Clinical Competence, were reported orally to all members of the Writing Committee at the first meeting, and were updated at each meeting or as changes occurred. Please see Appendix for the relationship with industry information pertinent to this document.
Mark A. Creager, MD, FACC, FAHA Chair, ACC/AHA/ACP Task Force on Clinical Competence and Training
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I. Introduction and background
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Both X-ray fluoroscopy and X-ray cinefluorography are core imaging techniques that make invasive cardiovascular procedures possible. Right heart catheterization for hemodynamic monitoring, diagnostic cardiac and vascular angiography, interventional cardiovascular procedures, clinical electrophysiologic studies, temporary and permanent pacemaker, and internal cardioverter-defibrillator placement are among the important cardiovascular procedures that either require or are facilitated by X-ray imaging. Although many patients derive great diagnostic and therapeutic benefit from these procedures, the use of ionizing X-radiationconstitutes an associated hazard that must be justified by the procedure's benefits.
In recent years the capability and complexity of invasive cardiovascular procedures have increased substantially. Originally, fluoroscopically guided procedures were principally diagnostic. Currently, many procedures are therapeutic as well. As procedures have become increasingly complex, they may employ greater fluoroscopic durations leading to the potential for greater patient radiation exposure. Refinements in radiologic equipment have improved image quality while reducing X-ray dose rates. However, even though technologic progress has reduced exposure rates, the greater exposure duration that attends more complex procedures may lead to an increased overall patient and operator exposure accompanied by a greater potential for radiation-induced injury. At present, although many patients derive great benefit from fluoroscopically guided procedures, some suffer radiation-induced injuries as an unintended consequence.
Radiation-induced patient injury takes many forms. Entrance port skin ulceration and necrosis are the most obvious and some of the most attention-getting. This complication consists of a localized area of non-healing skin necrosis at the site of X-ray beam entrance (typically on the patient's back). The shape of the affected area conforms to the shape of the X-ray beam entering the patient and may or may not have sharp borders. Examples are illustrated in Figures 1A to 1C (1–3). Entrance port skin ulceration causes considerable long-term morbidity often requiring skin grafting.
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Figure 1 (A) Radiation injury following angioplasty. This was the first reported case of radiation-induced skin necrosis from a coronary angioplasty procedure. The patient underwent three coronary angioplasty procedures, each of which lasted between 1 to 2 h. The last two procedures were performed on the same day, 6 months after the first, and involved about 1 h of fluoroscopy on-time. No data on the number of cine frames is available. Erythema was noted on the patient's back when the patient was removed from the table after the last procedure. One month later the patient reported erythema in the same areal; this persisted. The image on the left shows the appearance approximately 5 months after the procedures. The condition progressed into blistering, exudation, ulceration, and necrosis. The image on the right shows the wound 22 months after the third procedure (1). (B) Radiation injury from an electrophysiologic ablation procedure. A 52-year-old man underwent an ablative procedure for supraventricular arrhythmias. His arm had accidentally been positioned within the radiation field during the 10-h procedure. The estimated dose of radiation was in the range of 15 to 20 Gy (2). (C) Chronic radiation-induced skin injury. A 17-year-old girl underwent two electrophysiologic ablation procedures to treat an arrhythmia. The total fluoroscopy time was about 100 min. Erythema was present 12 h after the procedure. At 1 month, the area was red and blistering. At 2 years, the area was described as an atrophic indurated plaque with linear edges, hyperpigmentation and hypopigmentation, and telangiectasia. The patient had difficulty raising her right arm. Because of the close proximity of the breast to the X-ray beam, the scattered radiation resulted in a substantial dose to her right breast, placing her at an elevated risk for breast cancer (3).
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Complications associated with high doses of fluoroscopy were reported as a theoretical consideration, before they were actually observed and subsequently published in the refereed literature (4,5). By late 1994, the Food and Drug Administration (FDA) had received a sufficient number of reports of injuries to call attention to the problem with an advisory that later appeared on its Web site (http://www.fda.gov/cdrh/fluor.html). Although the first known modern-era event of an induced skin necrosis in the U.S. occurred in 1990, it was not reported in the medical literature until 1996 (1). Subsequent reports in the clinical literature documented that the problem is ongoing (6).
The injuries cited in the literature and to the FDA are likely only a small fraction of the total that actually occur. In addition, it may be suspected that there is the potential for X-ray–guided procedures to cause other less clearly attributable adverse effects, such as neoplasia.
The core principle governing the use of ionizing radiation is ALARA (As Low As Reasonably Achievable). The ALARA principle recognizes that there is no magnitude of radiation exposure that is known to be completely safe. This principle confers a responsibility on all physicians to minimize the radiation injury hazard to their patients, to their professional staff, and to themselves. To practice the ALARA principle, the physician must possess a basic knowledge base in two areas. He or she must know how to operate the X-ray fluoroscopic equipment in a manner that generates optimal image quality with minimal patient and clinical personnel exposure. The physician must also possess the knowledge to recognize patients and circumstances in which the risk of radiation-induced injury is increased. In these circumstances, the physician is responsible for considering that risk in case selection and in procedure conduct decisions.
To meet this responsibility effectively, physicians must possess an understanding of radiation physics, radiation biology, X-ray image formation, and the operation of an X-ray cinefluorographic unit. This knowledge base is well documented, and physicians are responsible for understanding it. Applying it appropriately in the interest of patient and clinical staff protection should be viewed as a standard of care.
This document's purpose is to serve as a resource to physicians who perform X-ray fluoroscopically guided procedures. Although not comprehensive, it contains an introduction to the field written specifically to be accessible and relevant to practitioners. It provides them with an introduction to and an overview of the requisite knowledge base needed to protect patients, clinical staff, and themselves from radiation-induced injury. Additional educational material is available in both the literature and in textbooks (7). It recommends a core curriculum that physicians who perform such procedures should study. The ACC previously published a document summarizing the radiation physics and biology knowledge base relevant to operator and clinical personnel radiation protection (8). This document is intended to be a companion to that document with a principal focus on patient protection.
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II. The physics and nature of X-radiation
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The nature of X-radiation.
X-rays are a form of electromagnetic radiation and have many characteristics similar to visible light as well as some important differences attributable to their greater energy content. X-rays are conveniently described in terms of photons—a quantum (discrete packet) of electromagnetic radiation containing a defined amount of energy. A stronger X-ray source produces more photons per unit time than does a weaker source.
An X-ray photon in the diagnostic range contains 5,000 to 75,000 times as much energy as a visible-light photon. A green light photon contains 2 electron volts (eV) of energy. The X-ray photons used for imaging have energies ranging between 10 keV (10,000 eV) and 150 keV. The difference in biologic effects between visible light and X-ray photons is largely attributable to energy content differences.
X-ray generation.
X-rays are produced in an X-ray tube when high-energy electrons, in an arc created by applying a voltage across a gap, interact with tungsten that forms the X-ray tube anode. (Details of X-ray tube construction can be found in the following text.) When a high-energy electron interacts with a target atom, a variable fraction of its energy is converted into an X-ray photon, and the remainder is dissipated as heat.
X-ray spectra.
X-rays generated by diagnostic machines contain a spectrum of photon energies that range up to a maximum determined by the voltage applied across the X-ray tube gap. The peak tube voltage (kVp) determines the maximum photon energy (expressed in kilo electron volts [keV]).
Mechanisms of X-ray absorption.
X-ray photons penetrate tissue to a varying degree. This phenomenon is the basis of X-ray imaging. The attenuation (the degree to which X-ray beam intensity decreases as it passes through an object) of any material is determined by four parameters:
- • X-ray photon energy
- • The atomic number of the atoms that make up the object
- • The physical density (g/cc) of the object
- • The thickness of the object
The penetrating power of an X-ray photon increases as photon energy increases. Low-energy photons have insufficient energy to penetrate tissue. Consequently, they do not contribute to image formation, but, as they are absorbed by the skin, they contribute to the patient's entrance port skin dose.
Several X-ray absorption processes occur at different photon energy ranges. Two X-ray absorption mechanisms that are important for diagnostic imaging are the photoelectric processand Compton scattering.
The photoelectric process occurs when a photon interacts with an orbital electron of an atom (generally the electrons in the K shell). In the photoelectric effect, the X-ray photon is completely absorbed and a free electron is ejected from the atom ionizing it. This process occurs preferentially at low photon energies with high atomic number absorbers. Photoelectric absorption is the principal absorption mechanism that renders iodinated contrast agent and metallic stents visible in X-ray fluoroscopy.
Compton scattering occurs when X-ray energies are much greater than the absorbing material's electron binding energy. The products of a Compton interaction are a scattered X-ray photon of lower energy than the incident photon and a recoil electron. The Compton affect is the main interaction process for diagnostic X-rays in tissue. Most stray radiation in a fluoroscopic laboratory arises from the Compton process.
X-ray dose.
Radiation dose is delivered by interactions of X-ray photons with the individual atoms that comprise a tissue. These interactions transfer energy from the X-ray photon to the atoms in the tissue. (The term "dose" as used here refers to the "absorbed dose.")"Dose" is a measure of the concentration of energy absorbed by tissue. The formal definition of dose is the amount of energy absorbed from the radiation field by a volume of tissue divided by its mass. Both the absorption of some radiation by the patient and the transmission of a sufficient quantity of radiation through the patient to the image receptor, are necessary to form a radiologic image. Image contrast is produced by partial absorption of the beam (Fig. 2). The dose delivered to a patient during an X-ray fluoroscopic examination is derived from the X-ray photons that enter but do not leave the patient. The most relevant of the many radiation measurement parameters and their relation to dose are shown in Table 1.
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Figure 2 Image contrast. Differential absorption of the X-ray beam by different parts of an object renders its internal structure visible in an X-ray transmission image. X-ray photons are represented as arrows. The left-hand example shows an object that allows all X-ray photons to pass through it unattenuated. In this circumstance, no image of the object is generated. The center example shows an object, different partsof which absorb different fractions of the incident photons modulating the beam intensity and generating an X-ray image of the object. The right-hand example is completely opaque to X-ray photons and absorbs all of them. The X-ray image of this object would be a silhouette with no definition of the object's internal structure.
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Measurements of radiation.
Radiation levels and doses can be characterized by utilizing a number of different parameters. The measurements are derived from a physical effect evoked by the radiation. The nomenclature has recently been revised. Currently, radiation quantities are expressed in SI (System Internationale) units and are listed and defined in Table 1. For comparison with the older literature, Table 1 also contains the corresponding earlier units shown in brackets (i.e., gray [rad]).
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III. Principles of X-ray image formation
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X-ray image generation.
The patient is illuminated with a beam of X-rays. Different structures in the patient absorb different fractions of the incident radiation, modulating the beam intensity (Fig. 2). The modulated beam that exits from the patient is detected by an image receptor. An object is delineated in an X-ray image if its X-ray absorbance is sufficiently different from that of its surrounding structures to produce a different exit beam intensity in that location. For example, an iodine-containing "contrast medium" employed to enhance the visibility of vessels absorbs significantly more of the radiation beam than does the blood it displaces, rendering the contrast-filled vessel visible on the X-ray image.
Parameters that affect X-ray image formation.
X-ray beam penetrating power
To produce an optimally exposed image, the X-ray beam's penetrating power must be appropriately adjusted for the patient's X-ray attenuation. This may be accomplished by varying a number of beam parameters. Ideal X-ray imaging parameters appropriately balance the requirement for contrast (needed to detect the object), the requirement for sharpness (needed to characterize it), and patient dose.
Increasing the kVp of an X-ray beam decreases its absorption, enabling the penetration of dense body parts, and reduces patient exposure by reducing the fraction of the beam absorbed by the patient. However, as kVp increases, the difference in the relative absorption of different tissues decreases. This decreases beam modulation and reduces image contrast. Therefore, optimal X-ray imaging requires a compromise kVp that produces the best balance of penetration power, image contrast, and patient dose.
Increasing the total electrical power applied to the X-ray tube, without changing kVp, can also increase the number of X-ray photons that penetrate the patient. This generates a greater number of X-ray photons of the same penetrating power. This strategy maintains image contrast at the cost of greater patient dose and greater X-ray tube loading. The gain from this strategy is an image with less noise, greater contrast, and better definition (see the discussion of image noise in the following text) at the cost of greater subject exposure and a greater X-ray tube loading. Another potential downside of this strategy is that the increased loading may require a larger X-ray tube focal spot, which will reduce image sharpness. In most modern cinefluorographic units, these parameters are set automatically by programs installed in the system. The programs are user-configurable if desired. It is important that physician users understand the operation and the patient exposure implications of choosing among the different selectable programs.
X-ray beam filtration
Because low-energy X-ray photons have very limited penetrating ability, they deposit their energy in a patient's superficial tissues, exposing these structures without contributing to image formation. Thus, it is desirable to remove (filter) low-energy photons from the X-ray beam. Aluminum has appreciable photoelectric absorption at low photon energies. Its attenuation decreases with increasing photon energy. Placing an aluminum disk on the output port of the X-ray tube preferentially removes low-energy photons from the beam, thus reducing the dose absorbed by the patient (Fig. 3). This process is called "hardening the beam." Increasing beam hardness increases the fraction of the beam's photons that successfully penetrate the patient and that contribute to the image. This means that less radiation must enter the patient in order to produce a given exit dose. Thus, if the beam is hardened, overall patient exposure is reduced even though the dose that reaches the image detector is the same. Many newer cardiovascular fluoroscopic systems use combinations of high-power X-ray tubes equipped with copper filters (producing a higher photoelectric absorption than aluminum) and special system regulation curves to produce even greater beam hardening while maintaining image contrast and quality. The effect of beam hardening on the distribution of X-ray photon energies is displayed in Figure 3.
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Figure 3 X-ray beam filtration. Plots of the X-ray photon energy composition of a typical beam of diagnostic X-ray produced at 70 kVp. The number of X-ray photons is plotted on the ordinate, and the photon energy is plotted on the abscissa. The original X-ray beam e (curve A) contains photons with multiple energies including many with energies below 30 keV that are of no diagnostic value. Intercepting the X-ray beam at the exit port with copper filters (curve B) preferentially absorbs low-energy photons. This reduces skin dose but also image contrast. Decreasing kVp to 60 while employing copper filtration (curve C) improves contrast but decreases beam intensity. Increasing the X-ray tube current increases the total number of photons produced and restoresbeam intensity (at the expense of additional X-ray tube loading) (curve D). The X-ray spectrum represented by curve D provides the same image quality as that provided by curve A at much a reduced dose to the patient.
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Scattered radiation
Scattered radiation is produced when the X-ray beam interacts with the patient. Scattered radiation that reaches the image receptor constitutes noise and reduces image contrast. Scattered radiation is also the principal source of exposure to both the patient' body parts that are outside the field of the primary X-ray beam and to the laboratory staff. The amount of scatter increases with increases in the size of the X-ray field and the intensity of the X-ray beam (Fig. 4). All measures that reduce patient dose commensurately reduce the dose of scattered radiation that exposes both the patient and the operator and clinical personnel.
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Figure 4 The intensity of scattered radiation as a function of exposed field size. For the same magnification mode of the image intensifier, scatter intensity increases everywhere in the room as the collimator is opened. Doubling the beam area doubles the scatter dose rates. The left-hand image illustrates the scatter field from a 60 left anterior obliqueprojection at a height of 100 cm from the floor with the collimator partially closed to allow exposure of only the central 50% of the field. The scattered radiation dose rates (which expose both the patient and the operator) are indicated by the isodose lines. The right-hand image illustrates an identical X-ray system position except that the collimator is fully open, exposing the entire field. Note that at any location, the intensity of the scattered radiation doubles. FOV = field of view.
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Image noise
A radiographic image of a uniformly dense object will have point-to-point variations in brightness. These random fluctuations, which are an inherent property of the X-ray beam, are called image noise. Noise is due to the quantum nature of the X-ray beam and increases as the X-ray dose decreases. Noise reduces the ability to detect low contrast structures. Increasing the dose suppresses noise and increases the ability to resolve structures. Image noise in an X-ray fluoroscopic image appears as a speckling of the image that is also commonly referred to as "quantum mottle." Image noise should be readily apparent in fluoroscopic images if X-ray equipment is properly calibrated. Fluoroscopic doses should be deliberately set at low levels to minimize total patient dose accumulation during the portion of the procedure that requires lesser image quality while reserving higher-dose (and lower noise) imaging for circumstances when image clarity becomes more critical. (It isimportant to point out that many current digital image processing algorithms are intended to decrease an image's apparent noise. Thus, the smooth appearance of an image acquired on a current digital system does not necessarily mean that the image was acquired at a high dose.)
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IV. The operation of an X-ray cinefluorographic unit
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Overview.
The main functions of an X-ray cinefluorographic system are to produce a collimated X-ray beam of appropriate intensity and quality, to project that beam through the patient at a desired angle, to detect the modulated X-ray beam after it passes through the patient, and to transduce the modulated X-ray beam into a usable visible light image. X-ray production is regulated by feedback loops from the image receptor. These components are schematically illustrated in Figure 5.
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Figure 5 Block diagram of a filmless X-ray cinefluorographic unit. The unit consists of a patient positioning system, an X-ray source, an X-ray image detector, and a digital video image processor, recorder and display system. X-rays are produced in the X-ray tube from highly controlled electrical power that is applied by the X-ray generator. X-rays that penetrate the patient form an X-ray image that is detected and converted to a visible light image by the image intensifier. The visible light image is detected by the video camera and converted to a digital video signal that is processed and displayed as a visible light image on video monitors. Feedback circuitry from the digital video processor communicates with the X-ray generator. This enables modulation of X-ray output to achieve appropriate subject penetration by the X-ray beam and, accordingly, proper image brightness. X-ray systems that have flat panel detectors rather than image intensifiers do not have video cameras as the flat-panel detector produces a digital video image directly without the intermediate visible light stage.
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X-ray generation.
The X-ray generator controls and delivers electrical power to the X-ray tube. It applies a high voltage across the gap between the X-ray tube cathode and anode and electrically heats the tube's filament. This causes the emission of electrons from the filament. The cathode current (expressed in milliamperes [mA]) determines the number of electrons liberated at the cathode. The voltage applied across the gap (expressed in kilovolts-peak [kVp]) accelerates the liberated electrons across the gap from the cathode filament to the anode and determines the energy with which they strike the anode material. The accelerated electrons interact with the metallic anode of the tube. A small portion of the energy carried by the electrons is transformed into X-rays. Thus, the cathode current (mA) determines the number of X-ray photons produced, and the tube voltage (kVp) determines the energy of the X-ray photons produced. The essential elements of a medium power X-ray tube are shown in Figure 6.
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Figure 6 Medium power rotating anode X-ray tube. The entire apparatus is enclosed in an evacuated glass envelope. Electrons are emitted from the cathode filament cathode cup. They are accelerated across the tube by the application of high-voltage (kVp) and impinge on the anode disk target. A small portion of the electron beam's energy is converted to X-rays upon impact with the anode. The majority of the energy heats the focal spot. Rotating the anode disperses the heat along a long track rather than concentrating it in a single spot. This enables the X-ray tube to generate a higher power beam.
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X-ray generation is inefficient from the standpoint of energy transformation. Less than 1% of the electrical energy applied to the tube is converted to X-rays; the remainder is deposited in the tube as heat. This creates an important heat dissipation challenge for X-ray tube design. Current tube designs are capable of dissipating several times as much heat as those of the early 1990s. Thus, these tubes can deliver significantly more radiation to patients without overload than was possible a decade or so ago.
For optimal imaging, the X-ray beam should emanate from an infinitesimally small point. This requires minimizing the size of the anode focal spot (the area on which the electron beam impinges) to as small a size as possible. However, the high power of the electron arc (approximately 100 kilowatts) limits the ability to reduce focal spot size because the power density at the focal spot would exceed the anode material's ability to absorb and dissipate the energy. As a result, the anode target would melt and destroy the tube. Therefore, at least two focal spot sizes are available on X-ray tubes: a large one, generally about 1 mm in size, and a small one of about 0.5 mm. The small size provides better image definition, but the X-ray output is limited and not sufficient for all tasks. The large size provides less definition, but permits a greater X-ray output when the task requires it.
Electrical current can be applied to the tube either continuously or in a pulse train (Fig. 7). This produces either continuous or pulsed X-ray output. Continuous irradiation causes images of moving objects to be blurred and leads to greater exposure. For this reason, virtually all present-day systems operate in a pulsed mode. The video frame rate is usually an exact multiple of the X-ray pulse rate in order to maintain synchrony between the X-ray pulses and the video acquisition. In the U.S., the standard video frame rate is 30 frames per second. Thus, typical X-ray pulse rates are 30, 15, and 7.5 pulses per second. The X-ray pulse rate determines the image sequence's temporal resolution. More rapid pulse rates provide greater temporal resolution and are useful for imaging rapidly moving structures, albeit at the price of a greater X-ray exposure. When the pulse rate is less than the video frame rate, the video chain presents the frame acquired during the pulse repeatedly (once for 15 pulses per second and three times for 7.5 pulses per second) until the next pulse is delivered. This eliminates flicker that would be caused by alternating bright and dark frames. However, it does not eliminate motion "jerkiness," which attends slower pulse rates. For most modern units used in pediatric cardiology, rates of up to 60 frames per second are available. These more rapid frame rates are needed to capture the rapid movement that occurs in small children's cardiovascular systems.
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Figure 7 Fluoroscopic pulsing X-rays are produced during a small portion of the video frame time (colored blocks). The pulse rate may be reduced by skipping video frames with the constraint that the video frame rate must be an exact multiple of the pulse rate to retain synchrony between the X-ray pulses and the video frame acquisition. In the U.S., the video frame rate is 30 frames per second. Thus, X-ray pulse rates of 30, 15, 7.5 frames per secondare feasible. The temporal relationships between X-ray pulsing and video framing are shown. Visual flicker would be apparent at pulse rates below 30 pulses per second. This phenomenon is reduced by repeatedly displaying the last image at the full video frame rate (30 frames per second). FPS = frames per second.
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Within the X-ray tube housing, X-rays leave the anode in all directions. The X-ray tube housing incorporates lead shielding that absorbs all X-rays except those emanating from its exit port. This defines the maximum diameter of the beam. In addition, a collimator is incorporated into the X-ray tube port to adjust the beam size to the minimum required for the imaging task. This beam collimation is necessary to confine the radiation to the imaged area of the patient, thus reducing exposure to other patient body parts. Collimation also reduces scatter radiation to clinical personnel.
X-ray image capture
Currently, two X-ray image capture systems are in active clinical use: image intensifier/video camera systems and flat-panel detectors.
Image intensifier-video camera systems
The X-ray image intensifier is a vacuum tube that converts the X-ray image into a visible light image that is brighter than that achievable by a simple fluorescent screen. Figure 8 is a schematic sagittal section of a single-mode X-ray image intensifier. The visible light output of the image intensifier is transmitted to a visible light image receptor such as a digital video camera for image display and recording.
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Figure 8 Diagram of an X-ray intensifier. The X-ray image passes through the input window and interacts in the input screen, converting the X-rays to a visible light image. The photocathode converts this light image to an electron beam image. The electrodesfocus and accelerate the electrons that strike the output screen, producing the amplified light image of a much reduced size.
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Flat-panel detector systems
The image intensifier/video camera combination is currently being superseded by integrated direct digital image receptors (flat panel detectors). These detectors incorporate a charge-coupled device visible light detector in direct contact with the input phosphor. Thus, they generate a direct digital video signal from the original visible light fluorescence without requiring the intervening stages. Figure 9 schematically illustrates the structure and operation of a typical flat-panel detector.
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Figure 9 Flat-panel detector (scintillator type). Panel A depicts a cross-sectional diagram of a flat panel detector showing its components. Panel B depicts a highly magnified view of a corner of the detector showing the individual charged-coupled array detector elements in contact with the cesium iodide scintillator and their connections to the readout electronics. The X-ray image interacts in the cesium iodide input screen, converting the X-rays to a visible light image. The input screen is in direct contact with a charged-coupled array detector that directly produces a digital output signal that is transmitted to the X-ray system video chain.
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X-ray exposure modulation.
The X-ray beam is attenuated as it passes through tissue. The degree of attenuation varies with tissue density and other factors such as the projection angle and the distance between the X-ray tube and the image receptor. In image intensifier systems, feedback circuits measure the brightness of the image generated by the image receptor. This feedback signal is used to modulate the output of the generator in response to changes in patient density and position. This is accomplished by an automatic dose rate control circuit designed to maintain a constant brightness level of the image intensifier output signal. X-ray intensity is increased if the detector measures a signal that is too dim and it is decreased if the signal is too bright. Similar feedback circuits are used in flat panel detector systems using the digital video output signal level as the source. Subject attenuation increases as overall tissue thickness increases. This means that the patient entrance port skin dose increases substantially when compound projection angles with cranial or caudal skewing are employed (Fig. 10).
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Figure 10 Effect of tissue thickness on entrance skin dose. Entrance skin dose increases rapidly with modest increases in the length of the radiation path in the patient. The causes are tissue attenuation and displacement of the patient's entrance surface toward the X-ray tube. Path length increases with compound beam angulation as well as with patient size. For a fixed geometry C-arm in this figure, the relative skin dose increases from 0.4 to 5.6 (14:1) as the path length increases from 12 to 36 cm (3:1). This example is for a mobile C-arm fluoroscope with fixed source-image distance and the patient is as close to the image receptor as possible. The actual relative values will differ for different types of equipment. Maximum fluoroscopic dose rates are limited by FDA regulations. There are no regulatory limits on cinefluorographic dose rates. Skin dose always accumulates far more rapidly in large patients (or with the use of compound angulation) than for non-compoundviews of thin patients.
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Video image capture.
Cameras and displays are the conduit between the image receptor and the observer's eye; they also serve to record the image permanently for later review and archiving. Various camera technologies have been used over the years. Initially, photographic film was the primary image-recording medium. However, the refinement of video cameras and digital image processing has led to the universal adoption of direct electronic video image recording. In image-intensifier–based systems, the amplified light image is monitored by a video camera. The video image is captured as a digital video image file. Flat-panel X-ray detector systems internally convert the X-ray image into a digital video signal. After processing, the image is displayed on a high-quality television monitor.
Image display and processing.
Digital images are usually processed before display or storage. Image processing alters the image with the intent of making it easier to interpret. All image-processing techniques involve compromise of some other aspect of image quality. Image-processing techniques include gray scale transformations to change contrast level; image smoothing to reduce the appearance of noise (at the expense of sharpness); and edge sharpening (at the expense of increasing the visibility of noise).
Image processing can also include combining multiple images. This reduces the effective noise level at the expense of blurring moving objects. Another image-processing technique involves subtracting one image from another to display differences between the two images (at the expense of increasing effective noise).
Cardiac cinefluorographic images are often acquired at a matrix size of 1024 x 1024 pixels. They are usually displayed in the laboratory at that resolution. Most cine archive systems (electronic or CD) downscan the images to a 512 x 512 matrix before storage. This is done to reduce data storage requirements and to reduce network data transmission times. Typically, downscanned images are re-upscanned to a 1024 x 1024 matrix size when they are displayed. This restores some but not all of the resolution of the original image. Thus, image resolution when viewed in the laboratory is somewhat better than when the image has been recalled from storage.
Imaging modes
The X-ray cinefluorographic units operate in two modes: fluoroscopy and acquisition (or image recording). The purposes and X-ray generator operating parameters of the two modes are different, particularly in terms of the input X-ray dose delivered. As a result, differences exist in image quality between these two modes. Figure 11 illustrates the differences in image noise between a single digital fluoroscopic frame and a single digital cine frame.
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Figure 11 Comparison of single fluoroscopic and fluorographic images. The left-hand image is a single frame acquired at a fluoroscopic level input dose. The right-hand image is a similar frame from the same patient recorded in the same projection, but acquired at an acquisition level input dose. Note that the left frame is noisier because a lower dose was used than for the acquisition frame.
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Fluoroscopy
Fluoroscopy provides a real time X-ray image when it is not necessary to record it. Fluoroscopy does not require the same level of image quality as does acquisition recording for diagnostic interpretation. Because these images are seen in motion, the neuropsychologyof vision effectively integrates several frames—effectively reducing the perceived image noise. Thus, greater image noise can be tolerated and fluoroscopic X-ray input doses can be lower than doses used for acquisition. As the fluoroscopic dose rate decreases, image noise increases, degrading image quality.
Fluoroscopic dose rates should be set at the lowest input dose rate needed to generate a usable image. Current fluoroscopic systems have two or more operator-selectable dose rates. The higher dose rates provide less image noise, thus enabling the delineation of greater detail at the cost of greater patient and operator exposure.
Acquisition (cine)
Acquisition mode generates higher resolution images suitable for diagnostic interpretation (including single-frame viewing) and archiving. Acquisition images are obtained at higher X-ray input doses in order to reduce image noise and optimize clinical visualization. Most X-ray cinefluorographic units are calibrated such that the per-frame dose for acquisition is approximately 15 times greater than for fluoroscopy. Thus, a single frame acquired in acquisition mode delivers about the same patient dose as 1 second of fluoroscopy. Figure 11 illustrates a comparison of single-frame images acquired at fluoroscopic and acquisition doses.
The optimal acquisition mode input dose per frame is that which achieves the best balance between image noise and image quality. The dose rate is also directly proportional to the acquisition frame rate. As with fluoroscopy, digital gap-fill can achieve flicker-free image displays at any frame rate. However, as frame rate decreases, image presentation becomes increasingly jerky despite gap-fill. As overall delivered dose is directly related to frame rate, the optimal frame rate is that which is just fast enough to capture clinically important transient events. A typical acquisition frame rate for adult studies is 15 frames per second.
Optimizing the exposure parameters that determine image quality.
The three main image-quality parameters—sharpness, contrast, and noise—are interdependent. The need to minimize patient exposure requires that dose be reduced to the minimum level that will generate an image with an acceptable degree of noise. The goal is to produce a usable image, not a perfect one.
For example, ideally an image would be acquired using a low kVp exposure to maximize image contrast and a large dose rate to minimize image noise. However, this would deliver a large patient exposure. Increasing kVp reduces patient exposure but decreases image contrast. Decreasing image receptor input dose reduces patient exposure but increases image noise. Thus, for a given patient density, there is an optimal compromise set of exposure parameters that preserve diagnostic quality image contrast at an acceptable image noise level while minimizing patient dose.
Fluoroscopy dose management issues.
Several features are available that facilitate patient dose reduction during fluoroscopy:
- • X-ray beam collimation: beam collimation restricts the size of the beam, enabling the operator to control the size and shape of the irradiated field. This provides an important exposure-limiting capability. The collimator should always be adjusted so that only the immediate field of interest is exposed to the X-ray beam, thus sparing the surrounding tissue from direct irradiation (see "Beam Collimation" in Section V).
- • Last image hold: This feature presents the last acquired fluoroscopic frame on the video monitor, thus providing an opportunity to study the image without continuing the exposure.
- • Pulsed fluoroscopy: This provides brief, several millisecond, X-ray pulses to generate images that are electronically processed by the digital video chain in order to furnish a continuously presented image. The pulse rate is operator selectable and ranges from 30 frames per second downward. As the pulse rate decreases, the patient exposure rate decreases (at the cost of increasingly jerky-motion presentation). Within limits, pulse rate reduction can produce images of acceptable quality for the purpose of the examination while minimizing dose rate. (Not all machines reduce dose rate with lower pulse rates. The dose-saving function of pulsed fluoroscopy can be assessed by a medical physicist.)
- • Virtual collimator and semitransparent diaphragm control: Current X-ray units have the ability to generate a virtual display of collimator and semitransparent diaphragm positions. This enables the operator to position these devices as desired without using fluoroscopy.
- • X-ray stand position memory: Current units are able to store multiple stand positions in memory and can automatically move to a selected position on command. This enables the operator to avoid the use of fluoroscopy to achieve a desired stand position.
There is a limit to which dose and frame rates can be reduced in cardiovascular applications. The neuropsychology of vision provides a degree of integration that decreases perceived visual image noise and jerkiness. In addition, digital image processing permits the digital equivalent of integration by a process called "recursive filtering." Application of recursive filters reduces the impression of noise at the expense of blurring objects that are moving. Fifteen frames per second is an optimal compromise frame rate. Rates below 15 frames per second may degrade the image presentation sufficiently to interfere with intricate device manipulation procedures, but may be adequate for less demanding tasks.
The dose rate often increases as the degree of electronic magnification of the image increases. For an image intensifier, the dose rate is roughly inversely proportional to the input area of the image detector. This relationship was obligatory for older film-based X-ray units as image-intensifier light output was directly related to input field size. For example, if the acquisition mode input dose for a 23-cm image intensifier is 100 nanograys (nGy)per pulse, the corresponding input doses when that intensifier is operated in the 17-cm and 13-cm modes are 183 and 313 nGy per pulse, respectively. For current fully digital image intensifier units, this relationship is no longer necessarily true as the light requirements of the video chain can be changed to require different amounts of light input at different image-intensifier field sizes. Similarly, the relationship between dose rate and active field-of-view for a flat-panel detector can be adjusted for different input field sizes. In general, both image-intensifier detector and flat panel detector dose rates are programmed to increase somewhat as the size of the field-of-view decreases, but these dose increases are required principally to reduce the image noise increase that would otherwise be noticeable at greater degrees of magnification (Fig. 11).
Acquisition dose management issues.
The acquisition mode is employed when images need to be reviewed and archived, and analysis of a static single frame or series of frames is necessary. The optimal acquisition mode input dose per pulse is that which achieves the best balance between image noise and patient dose. When moving objects are imaged, image sharpness is also determined by pulse width, with briefer pulses yielding greater sharpness. As with fluoroscopy, digital gap-fill can achieve flicker-free image displays at any acquisition pulse rate. The dose rate is also directly proportional to the pulse rate. The optimal pulse rate is that which is just fast enough to capture the transient moving events being examined. Thus, for adults it is now common to acquire images at 15 frames per second. Higher pulse rates are generally needed to image small children. Substantially lower pulse rates are usually appropriate for the peripheral vasculature. Image sharpness is related to the pulse width not the frame rate.
Digital image subtraction.
A digital subtraction image is obtained by subtracting one image from another. This electronically removes information that is identical in two images. The resulting image is a display of the difference between the two images. In angiography, the first image (mask) is obtained before the injection of contrast. The second image is acquired during the angiographic run. The computed resultant image contains the difference between the two acquired images and emphasizes the contrast-opacified structures. Because the subtraction process accentuates image noise, it is necessary to counter this effect by acquiring each of the original images at a substantially (as much as 20-fold) higher dose per frame. The increased dose per frame may be partially offset by the ability to employ slower frame rates. However, studies that use digital subtraction imaging generally employ larger aggregate doses than do studies that employ unsubtracted cinefluorography.
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V. Determinants of patient X-ray dose
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Patients undergoing invasive cardiac procedures do not receive uniformly distributed whole-body radiation. The delivered radiation is mostly concentrated in a confined area of the thorax. The effect of the patient's exposure is related to the dose received by each directly or indirectly exposed structure.
Measurements of patient dose.
Two parameters of dose—the dose at the interventional reference point (IRP) and the dose-area product (DAP)—are useful for characterizing patient exposure. Currently available interventional fluoroscopic equipment determines real-time estimates of the instantaneous and cumulated values for these dose factors. The unit's indication of these cumulated values provides valid indicators of a patient's dose and consequent risk for radiation-induced effects.
Dose (air kerma) at the IRP.
The IRP is located on the X-ray beam axis 15 cm from isocenter on the X-ray tube side (Fig. 12). For an interventional cardiologic procedure, this location approximates the location of the skin at the beam entrance point when the heart is located at isocenter. Thus, the air kerma at this point is an indicator of skin dose. Although quantitatively not an exact measure of skin dose, the estimate provides a measure by which an assessment can be rendered regarding the risk of injury to the patient's skin.
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Figure 12 The interventional reference point (IRP). The IRP is located 15 cm from isocenter on the X-ray tube side of a fluoroscope, as shown in the diagram. This approximates the location of the patient's skin in cardiologic procedures. As is obvious, for conditions other than the one illustrated here, the IRP might be located many centimeters away from the skin surface. This might occur for non-isocenter settings, larger or smaller patients, or different beam orientations. Therefore, air kerma measurements at the IRP must be used for guidance purposes and not considered to be true skin dose.
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DAP.
The DAP is the absorbed dose to air (air kerma) multiplied by the X-ray beam cross-sectional area at the point of measurement. It is expressed in Gy·cm2 or some variation thereof. The cumulated DAP for a procedure is a surrogate measurement for the total amount of X-ray energy delivered to the patient. Consequently, it is a measure of the patient's risk of a stochastic effect (see Section VI).
Contrary to the measurement of dose at the IRP, the value of the DAP of an unattenuated X-ray beam does not depend on the distance from the X-ray source. This is because, although the dose decreases with distance from the X-ray source, the beam area increases commensurately. The DAP is usually measured by means of a transmission ionization chamber placed in the X-ray tube assembly. It may also be calculated from generator and collimator settings. A typical beam cross-sectional area at the IRP is between 30 and 100 cm2. Thus, the DAP at the IRP might numerically be 100 to 300 times the air dose at the IRP.
Value of dose monitoring.
Estimated values for air kerma at the IRP are calculated and displayed by modern X-ray units. Typically, the system measures the DAP using an ionization chamber placed in the X-ray tube's exit port. As discussed previously, this parameter is constant at all distances from the beam. The system calculates the dose at the IRP from DAP data and the known position of the X-ray tube collimator leaves.
However, the values that the system displays, nonetheless, are estimates and have a margin of error that may be as much as a factor of two or greater. Thus, the measure, therefore, must be interpreted with some discretion. For example, a calculated IRP air-kerma dose of 2 Gy is very unlikely to represent an actual skin dose of 6 Gy (the approximate threshold dose for delayed skin erythema). Conversely, if the calculated IRP dose is 4 Gy, it is more likely that the erythema threshold may have been crossed. Thus, patients who receive an IRP dose greater than 4 Gy should be advised that they might develop a skin rash, along with instructions on what to do in the event that one is observed.
Dose levels at the IRP and DAP are influenced by many variables, not all of which are under the operator's control. Nonetheless, assessment of these parameters provides a measure of a physician's radiation management performance. Factors not under the operator's control include patient size and disease complexity. However, other variables, such as X-ray system position, collimator position, and appropriateness of beam-on time, are affected by the operator's attention to radiation safety practices. Thus, although the relationship of DAP to patient injury is indirect, monitoring DAP is a valuable part of overall quality assurance monitoring. The DAP tracking for all procedures provides a measure of appropriateness of patient radiation protection practices.
Dose at the IRP monitoring.
The dose at the IRP has a direct relationship to the risk of patient skin injury. Real-time intraprocedural monitoring of the dose at the IRP may be employed to make decisions about procedure conduct. If a large dose has been delivered before completion of non-critical aspects of a procedure, it may be appropriate to abbreviate that procedure with plans to reassess the patient's condition at a later date. This would permit skin examination to determin |
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