Fluoroscopy systems provide real-time or near real-time x-ray imaging of patients. Fluoroscopy systems provide the temporal resolution necessary for the operator to use image guidance for the placement of medical devices (e.g., catheters, stents, etc.) or to observe temporal physiological phenomena. Common uses for fluoroscopy include upper and lower gastrointestinal studies, where the radiologist may manipulate barium or air contrast around the GI tract, in order to observe bowel obstruction, anatomical abnormalities, or polyps. The fluoroscopy imaging system is also a crucial tool in the placement of catheters and devices for vascular imaging and interventions, such as for abdominal, cerebral, or cardiac procedures. Other uses of fluoroscopy include percutaneous needle biopsy, swallowing studies, hardware placement and positioning during surgery, etc.
There are two basic modes of operation for fluoroscopic systems: (1) fluoroscopy, which provides real-time imaging for positioning, which is generally not recorded and is relatively low in radiation dose, and (2) fluorography, which essentially uses the fluoroscopic imaging chain in a pulsed radiographic mode to record and document clinically relevant temporal sequences such as blood flow through vessels (angiography), mechanical motion of joints, etc., with higher levels of radiation consistent with radiography.
Image intensifiers were the central technology enabling low-dose fluoroscopy since the early 1960s. These earlier fluoroscopy systems used continuous x-ray sources with 30 frames per second (FPS) acquisition rates, analog TV cameras for fluoroscopic viewing, and film-based devices for recording. Modern fluoroscopic systems make use of pixelated flat-panel detectors (FPDs) with large, multiple fields of view (FOVs), numerous modes of operation, a pulsed x-ray beam to reduce vascular motion, flexible frame rates, and numerous image processing techniques such as digital subtraction angiography (DSA) and road-mapping, and some systems even have cone beam CT acquisition capabilities.
The radiation dose rate for fluoroscopy has been reduced substantially with improvements in x-ray tube technology, replacing minimally filtered (Al) high tube potential (90 kV) imaging of the past with highly filtered (Cu), low x-ray tube potential (e.g., 60 kV) imaging. With these improvements in the x-ray spectra used in fluoroscopy, not only have the dose rates been reduced considerably, but image quality has improved as well. More complicated fluoroscopic suites include numerous methods for radiation safety of the staff, including ceiling-hung clear x-ray shields for protecting the upper body of the operator, along with numerous other devices deployed for reducing the scattered radiation received by staff.
The modern fluoroscopy system embraces both fluoroscopy viewing and fluorography recording and many imaging protocols designed for specific imaging applications. All modern systems embrace sophisticated computer interfaces, large ceiling-hung display monitors, and tableside controls enabling dozens of operational modes. All these features challenge the physician operator to become familiar with the increasing complexity of these systems and learn the most appropriate protocol
for optimizing the acquisition of patient information while reducing radiation dose levels to both the patient and staff.
Fluoroscopic systems are flexible general-purpose devices that are configurable to meet the requirements of a wide range of clinical procedures. Most fluoroscopes have hundreds of preprogrammed imaging protocols intended to optimize many more procedures than might be obvious from the system’s “type.” Image acquisition and display controls include x-ray generator settings including automatic exposure rate control (AERC), image-receptor FOV, image processing methods, and image display adjustments. Clinically appropriate combinations are combined into “Examination Protocol Selection Buttons” (EPSB). Operators should understand these controls as well as the other aspects of the imaging system as they are a central node in the fluoroscope’s control loops (Fig. 9-1
▪ FIGURE 9-1 Block diagram of an interventional fluoroscope. Key components and control loops are shown. The operator is a key control element with the ability to select and modify radiation factors, imaging geometry, image processing, etc. (© Stephen Balter.)
9.1 FLUOROSCOPIC IMAGING CHAIN OVERVIEW
The principal feature of the imaging chain that distinguishes fluoroscopy from radiography is the fluoroscope’s ability to acquire acceptable real-time x-ray images with potentially high frame rates and lower dose per image. An up-to-date fluoroscopic imaging chain is shown in Figure 9-2
. Its key component is the solid-state image receptor, commonly called a flat panel detector (FPD). This device converts the x-ray signal into a series of digital images.
Digital radiography (DR) systems use a similar image receptor. Older DR systems acquired single images, and current advanced systems can acquire low frame-rate sequences. The underlying technologies are close to convergence, and future systems with tunable gain capabilities may provide both fluoroscopy and radiography with the same hardware.
The x-ray tube in an angiography system with substantial copper beam filtration combined with low kV allows angiography systems to achieve lower patient radiation
dose compared to older systems, while still delivering high iodine image contrast for angiographic applications. A collimator changes the size of the x-ray beam in response to the operator limiting the active FOV to a clinically relevant area, to adjust for changes in the source-to-image-receptor distance (SID) or when the overall FOV of the image-receptor is changed, commonly due to the operator changing the magnification mode.
▪ FIGURE 9-2 Diagram of a fluoroscopic system. The fluoroscopic imaging chain is illustrated, with the patient in the supine position. This figure includes a flat panel detector system. The x-ray system includes a collimator with motorized blades that automatically adjust to conform to the current field of view (FOV) and source-to-image-receptor distance (SID).
The basic product of a fluoroscopic imaging system is a projection x-ray image similar to a radiograph; however, a 20-min “on time” interventional fluoroscopic procedure, conducted at 15 FPS, produces a total of 18,000 individual fluoroscopic images, in addition to a potentially large number of fluorographic (e.g., digital angiography [DA], DSA, cine) images. Due to the large number of images (also referred to as frames), fluoroscopic systems must produce each frame with much less dose than a comparable radiograph.
Fluoroscopes use sensitive, low noise detectors. Standard fluoroscopy typically uses a 10-100 nGy detector dose per image, the amount often dependent on frame rate. Fluorographic images use 100-1,000 nGy per image. DR systems operate at 1,000-5,000 nGy air kerma to detector per image, and computed-radiography detectors use a range from 3,000-10,000 nGy per image.
9.3 FLUOROSCOPIC X-RAY SOURCE ASSEMBLY
9.3.1 X-ray Tube
The general construction of all x-ray tubes is similar. However, due to the nature of the procedures, fluoroscopic x-ray tubes are commonly used in a more continuous manner than radiographic tubes. These tubes also operate at radiographic power levels during fluorography. Fluoroscopic tubes have higher heat-storage capacity and faster cooling rates than radiographic tubes. Traditionally, the small focal spot is used for fluoroscopy and the large focal spot for fluorography. Thus, spatial resolution is likely to decrease when going from fluoroscopy to fluorography. Some tubes also have an additional micro-focus focal spot intended for geometrically magnified procedures. Newer systems select the smallest available focal spot that can accommodate the immediate x-ray power level for all modes of operation. This will result in similar spatial resolution for thin body parts and lower resolution for fluorography of thick parts.
Fluoroscopic collimators are designed to dynamically adjust the overall x-ray field size in response to a combination of inputs from both the operator and the system. The intention is to confine the x-ray beam to the smallest area within the active area of the image receptor that is consistent with immediate clinical requirements. Many systems also have additional, operator controlled, x-ray translucent elements (sometimes called wedge filters) that are used to equalize image receptor inputs when anatomical structures providing very different attenuation(e.g., lung and mediastinum) are simultaneously in the beam.
Portions of the x-ray beam that are not seen by the imaging chain do not supply any clinical information. Such irradiation unnecessarily increases both patient and staff radiation risk. Scatter from these areas also degrades image quality in the useful image.
Confinement of the beam to the active FOV is a system function. For both the image intensifier and the flat panel, the x-ray collimator automatically adjusts to limit the x-ray beam to the active FOV whenever the FOV is changed and to accommodate changes in the source-to-image-receptor-distance (SID). Ideally, the system should be set so that an unirradiated area is always seen around the edges of the image.
Collimation of the beam within the active FOV is the operator’s responsibility. Smaller beams produced by collimator adjustment by the operator reduce total
irradiation, thus improving both safety and image quality. The goal is to only irradiate the region of immediate clinical interest. Obtaining a smaller FOV by selection of a magnification mode does reduce the beam size and magnifies the displayed image; however, most fluoroscopes increase the dose-rate when the FOV is reduced by this method. Depending on programming, this increase may diminish or even negate the radiation saving effects of collimation within the FOV. Image magnification of a collimated area within a larger FOV can also be offered by the image processor through pixel replication and interpolation, or by simply viewing the image on a larger monitor. If the resulting image is clinically acceptable, this can reduce radiation dose relative to the use of magnification mode.
9.3.3 X-ray Spectral Shaping Filters
By 1990, the DQE of the image intensifier was near 50% and x-ray quantum noise was the major noise source in fluoroscopy. There was only a factor of two available for reducing patient irradiation by improving the efficiency of the image receptor. Radiation levels can also be reduced when the radiological conspicuity of clinically important items, such as contrast media and guidewires, is increased.
One way of improving the conspicuity of higher atomic number (Z
) elements is to increase the fraction of x-ray photons in the beam with energies slightly above the K absorption edge of the material of interest (e.g.
, iodine at 33.2 keV). Spectral shaping can be accomplished by adding a copper filter (0.1-1.0 mm) and simultaneously restricting the x-ray tube’s operating kV (Fig. 9-7
). Adding copper reduces the number of photons in the beam below the iodine edge. Reducing the kV limits the number of higher energy photons. For a given amount of electrical power consumed by the tube, such kV-filter combinations decrease the photon fluence available to form the image. Adequate fluence using spectrally shaped beams became feasible when x-ray tubes were developed with high power ratings.
Most modern fluoroscopic systems include variable thickness copper filters in the collimator assembly. The thickness of copper used at any given moment is automatically controlled by the AERC and operator selection of settings. In some systems, the acquisition selection (e.g.
, low-dose-rate fluoro, standard-dose-rate fluoro) determines the filter thickness. In other systems, the AERC further controls filter thickness
along with the usual x-ray factors. In some cases, for increasing patient thickness or acquisition mode changes, the limits of the x-ray tube power output require the filter thickness to decrease and the kV to increase. In such settings, at the same kV, the half value layer (HVL) will differ between fluoroscopy and fluorography; and the HVL may decrease with an increase in kV if there is a simultaneous decrease in copper thickness. Figure 9-8
illustrates this behavior for a typical interventional fluoroscope.
▪ FIGURE 9-7 Shaping the spectrum of the x-ray beam. A. The x-ray spectrum emerging from the tube is modified by a copper layer. The overall intensity of the beam is reduced, and the resultant spectrum has shifted to higher average energy. This figure shows the original spectrum, the attenuation coefficient of iodine as a function of energy, and the resultant modified spectrum with a higher fraction of photons above the iodine K edge. B. Increasing kV shifts the spectrum further, but most of the higher energy photons are far from the iodine K edge and do not contribute very much to iodine visualization. Tungsten K characteristic photons are seen here as well. C. Reducing kV moves more of the spectrum toward the iodine K edge. There are fewer low energy photons that would contribute to skin dose but not to the image. The photons, with energies above 60 kV shown in B are not produced. X-ray tubes need to have high power capabilities so that they can provide adequate x-ray flux without increasing kV. (© Stephen Balter.)
▪ FIGURE 9-8 AERC controlled Cu spectral shaping filter thickness. Measured data from interventional fluoroscope installed in 2010. Layers of poly methyl methacrylate (PMMA) and lead were used to simulate different patient sizes. The AERC selected a kV and filter thickness based on the image-receptor signal. Separate algorithms were used for fluoroscopy and fluorography (cine). Note that although the kV was similar at every step, the filter used for cine declined much more rapidly than for fluoro. The removal of copper is the reason for a higher cine HVL at 80 kV than 90 kV. (© Stephen Balter.)
9.4.1 Automatic Exposure Rate Control
Fluoroscopes offer at least two fluoroscopic and one fluorographic sub-modes. Often, systems offer literally dozens to hundreds of sub-modes to accommodate a wide range of operator preferences and clinical requirements. Most of these use an AERC system. The purpose of AERC is to deliver a constant x-ray intensity to the image receptor irrespective of tissue thickness in the beam path. This results in a reasonably constant detector SNR over its working range. In a FPD, the AERC collects detected x-ray intensity information observed by a predetermined subset of dexels. Depending on clinical configuration, the subset ranges from dexels in a relatively small central area to all dexels within the collimated beam. If the average signal is too low, the AERC commands the generator to increase x-ray output. Similarly, too high a signal initiates a command to reduce radiation output. AERCs commonly further adjust outputs when the FOV or frame rate is changed. These adjustments stabilize the image noise perceived by the fluoroscopist when these factors vary.
For large patients or steep projection angles, the x-ray air kerma rate may reach the normal fluoroscopy regulatory limit (88 mGy/min). The resulting x-ray flux
reaching the image receptor (II or FP) is lower than desired, and x-ray quantum noise is higher than desired. Necessary additional electronic amplification may add additional noise. The noise in the resulting image may be too high for some clinical purposes. Some fluoroscopes have a specially activated high-level control (HLC), which enables a higher limit (176 mGy/min) to reduce noise. The high-level mode should be only used when clinically necessary.
Under AERC, the system changes the x-ray factors (kV, mA, pulse width), focal spot, and beam filtration in a predetermined manner that varies between systems. As noted above, the same system usually offers user-selectable options with markedly different behaviors. These selections are part of the toolkit used to select the balance between patient dose and image quality appropriate to different types of examinations, and in some cases different operator imaging requirements. The change in output, and the maximum output dose rate when changing from “low” to “normal” to “high-level” at any patient thickness is also included in the system’s configuration. Image processing parameters may also be controlled by the AERC. These changes further influence image presentation to the operator.
illustrates some possibilities with a hypothetical system equipped with three fluoroscopic modes, each with a different regulatory exposure-rate limit. In Figure 9-9A
, the output is the same until the mode’s limit is exceeded. All three modes produce the same radiation output for thin to medium size patients. The “low” mode is limited to an air kerma rate (for C-arm gantries, at 30 cm from the image receptor assembly) of 44 mGy/min; in this mode all thicker patients receive the limiting dose rate while image quality decreases as patient thickness increases. Similar behavior is seen for the “normal” and “high-level” modes when their limits are reached. Figure 9-9B
illustrates a system with three fluoroscopic modes. The lower two fluoroscopic modes converge near the regulatory limit. The high-dose-rate mode is higher for all patient thicknesses and has double the usual regulatory limit.
illustrates that the relationship between patient thickness and patient dose for any mode and dose rate limit is programmable. The upper curve is intended to maximize the visibility of iodinated contrast media as much as possible. The generator is programmed to keep the kV low and adjust mA and/or pulse width to provide an appropriate detector signal. The lowest curve is intended to minimize patient irradiation at the possible expense of iodine visualization. It does so by rapidly increasing kV with increasing patient thickness. One or more additional intermediate curves may be provided on a given system.
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