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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_5196_Библиотеки_им_академика_М_И_Перельмана.pdf
CHAPTER 11 Equipment Design for Radiation Protection
207
• Digital radiography (DR)
• Digital fluoroscopy (DF)
• Digital mammography (DM)
Digital Imaging Overview
All non-analog or digital images have many similar properties regardless of the engineering technology with
which they are acquired, simply because they are all numerical approximations of non-digital (analog) signals.
Even though these signals may be produced by vast numbers of x-ray photons, the overall strength of the signal is
recorded as one number at each location on the image
receptor. All electronic radiography devices have inherent
limitations with respect to both spatial and contrast resolution due to the finite dimensions of their detector elements or picture elements (pixels). All digital imaging
devices are also subject to certain artifacts (i.e., effects
seen in an image system that were introduced by the technology used in acquiring the image or images) such as
aliasing, moiré patterns, and contouring (see Box 11.3).
These unwanted effects occur because digital images
are produced collectively by an array or matrix of elements that have finite dimensions and are always subject
to random quantum noise effects (mottle), which grow
in importance as the matrix elements become finer.
With any of these modes of computer-processed radiography, the latent (unprocessed) image formed by x-ray
photons on a radiation detector is an electronic image.6
Because this anatomic information is subsequently collected by a computer and shown on its display, it is called
a digital image.7 The familiar radiographic densities then
appear as levels of brightness associated with shades of
gray. The shades of gray that are displayed on a computer
screen constitute the contrast in the image. The number
of different shades of gray that can be stored in memory
and subsequently displayed is known as the grayscale. All
digital images are composed of numerical data that can
be easily manipulated by a computer.
The numeric values of the digital image are aligned in
a fixed number of rows and columns (an array) that form
many individual miniature square boxes, each of which
corresponds to a particular place in the image. These individual boxes collectively constitute the image matrix.
Each miniature square box in this matrix is a picture ele-
ment, or pixel. The pixels collectively produce a two-
dimensional representation of the information contained
in a volume of tissue.8 The size of the pixels determines
7
BOX 11.3 Image Aliasing, Morié
Patterns, and Contouring Defects
In computer graphics, aliasing is an image distortion
that shows up most simply as the jagged, or sawtoothed, appearance of curved or diagonal lines.
Moiré patterns occur when an object that is being imaged contains many fine, repetitive details. As a result,
strange-looking wavy patterns are overlayed on the image. This is shown in the figure below.
(From Carter CE, Vealé BL: Digital radiography and PACS,
ed 3, Philadelphia, 2019, Elsevier.)
Contouring artifacts usually show up as patterns of
small blocks in an otherwise smooth image. An example of this are adjacent groups of rectangular or blocklike images of different intensity or brightness superimposed on what is really a clear smooth sky. These
blocks become very much more noticeable as the image is viewed on larger devices or monitors.
the sharpness of the image. The resolution is finer when
pixels are smaller. One example of a standard matrix size
is 512 pixels high by 512 pixels wide, or simply 512 3 512.
A range of other matrix sizes are also used, such as 1024
3 1024. For the same field of view (FOV), the latter corresponds to four times as many elements distributed over
the same area. The pixels are therefore smaller, which
208
CHAPTER 11 Equipment Design for Radiation Protection
Indirect Conversion
Scintillator
Photodiode
Thin-film
transistor array
Electrical signals
Fig. 11.13 Some large area detectors provide indirect conversion of x-ray energy to electrical charge through
intermediate steps involving photodiodes or charge-coupled devices. Other area detectors provide direct
conversion of x-ray energy to electrical charge through the use of a photoconductor. (From Hendee WR, Ritenour ER: Medical imaging physics, ed 4, Chicago, 2002, John Wiley & Sons.)
Scintillator
Charge-coupled-
device array
Electrical signals
leads to improved patient image detail. It should be noted
that such resolution increases, however, are associated
with very substantial increases in the quantity of data to
be computer processed.
Digital image receptors used in DR convert the energy of x-rays into electrical signals. The image receptor
is divided into small detector elements that make up the
pixels of the digital image. There are various types of
these image receptors. Some use a scintillator, such as
amorphous silicon,* to convert the x-ray energy into
visible light. The visible light is then transformed into
electrical signals by an assortment of transistors or an
array of charge-couple devices (CCDs), such as those
found in video cameras. Other systems use a photocon-
ductor, such as amorphous selenium,* to convert the
x-ray energy directly into electrical signals that are then
read by an ordered grouping (matrix) of transistors. In
these systems, the number and size of small transistors
or CCDs determine the number and size of pixels in the
digital image. Advances in materials technology have
resulted in pixel sizes as small as 50 micrometers, which
approaches the resolution of screen-film imaging systems (Fig. 11.13).
*Amorphous refers to a non-crystalline grouping of silicon
atoms (
in a regular geometric pattern, the silicone atoms or the selenium atoms are distributed in a continuous random fashion.
Si28) or selenium atoms (34Se79) in which, rather than
14
Direct Conversion
Photoconductor
Thin-film
transistor array
Electrical signals
Digital images can quickly be accessed via a PACS*
network at multiple workstations at the same time,
thus allowing image viewing to be very convenient
for physicians providing patient care. Patient information and reports can be included in the patient’s
imaging file, along with records from other imaging
modalities.
9
Computed Radiography (CR)
Computed radiography CR is the descriptive term for
those x-ray systems that generate images using the
process of photostimulable luminescence (PSL). In this
technique, the energy of the conventional diagnostic
x-ray beams passing through a patient is not projected
onto x-ray film within a cassette but rather is deposited
onto a modified crystalline material known as a photostimulable phosphor (PSP).** The absorbed energy
*PACS refers to a picture archiving and communication system
imaging technology which provides throughout an authorized
and secure computer network both economical storage and convenient remote access to images from multiple modalities. The
universal format for PACS image storage and transfer is called
dicom (Digital Imaging and Communications in Medicine.)
**Europium-activated barium fluorohalide (BaFX:Eu) is the
most commonly employed phosphor in which “X” represents a
combination of a bromide (e.g., potassium bromide KBr) and an
iodide (e.g., sodium iodide NaI), typically 85% and 15%, respec-
1
tively.
CHAPTER 11 Equipment Design for Radiation Protection
209
produces a latent (undeveloped) image composed of
PSP molecules with electrons that have been excited
(energy raised) into fixed states, which are also known
as local potential energy traps.
The number and locations of these trapped excited
electrons is a molecular photo of the originally projected x-ray energy. The electrons are released from
their traps by gaining added energy from absorbing
external light of a specific wavelength (photostimulation). These electrons subsequently lose their excess
energy by the release of the photostimulable light (PSL)
they absorbed. An image-reading unit employing a
helium-neon laser is used to scan the light, which then
can be measured and enumerated to create a digital image. A computer stores the digital image for visual display on a monitor. If desired, the image can be printed
on a laser film when a hard copy is needed. While the
digital image is displayed on a monitor, the radiographer, by manipulating the computer mouse (Fig. 11.14)
or scrolling a touch screen cursor, can adjust the image
to the correct:
• Size
• Brightness
• Contrast
After adjustments have been completed, the image or
images are sent through a PACS system for review by a
radiologist. In CR, the receptor can be erased with white
light and reused to acquire another projection.
CR was useful as an interim step in converting older
imaging equipment to digital techniques, since the CR
plate or cassette could be substituted for the film-screen
cassette in older equipment with no further modifications. However, the speed of image acquisition and dose
efficiency considerations indicate that there will be an
eventual conversion of all equipment to DR in the future.
Kilovoltage. The kVp controls the penetrating ability of
the x-ray beam as it passes through human anatomy
and also affects radiographic contrast in the image. The
size of the part or area of the body to be imaged and the
type of subject contrast desired in the completed image
determine kVp selection.2 In CR imaging, unlike filmscreen imaging, it has been found that with respect to
kilovoltage, there is a substantial amount of flexibility
available for selection of the degree of desired subject
contrast. Regulatory standards require that technique
charts indicating optimal kVp values for all CR projections must be available in the x-ray room near the operating console for radiographers.
X-Ray Beam Collimation. For the computer to form a
CR image correctly, the body area or part being irradiated
must be accurately positioned in or near the center of the
CR image receptor. In practical application, only one
projection per image is obtained on a CR imaging plate.
Fig. 11.14 The radiographer at the monitor uses the mouse to
adjust the computed radiography image of the body part to the
proper size, density, and contrast before electronically sending
the image for reading.
Use of Radiographic Grids. The CR imaging plate can
absorb low-energy scattered photons; therefore, it is sensitive to scatter radiation both before and after it is sensitized by exposure to a radiographic beam.3 Because
of this increased sensitivity, a radiographic grid should be
used more frequently. For chest radiography, Carlton and
Adler advocated the use of a grid for optimum images
when chest measurements exceed 24 to 26 cm.3 Some CR
imaging manufacturers recommended the use of a grid
for particular radiographic projections that require relatively high-kVp settings. Grid selection depends on several factors: the size of the anatomic features to be
imaged, kVp selected, amount of scatter removal preferred, and grid frequency (lines per centimeter or inch).
It is customary and, for the best image quality, necessary to use a grid for anatomy sections more than 10 cm
thick or for techniques that exceed 70 kVp. This need
remains true with both CR and DR. In general, the
3
210
CHAPTER 11 Equipment Design for Radiation Protection
problem a radiographer faces with CR is that the mAs
required and, consequently, the patient dose received
usually are higher than for non-CR. The addition of a
grid will only further increase that dose. Many quality
assurance teams, however, realize that CR, because of
its wider exposure latitude, reduces the need for grid use
on the pediatric population. As a result, satisfactory
non-grid pediatric protocols have been developed and
implemented.
Digital Radiography (DR)
DR systems, unlike CR, use as image receptors activematrix flat panels consisting of a detection layer deposited
over an array of thin-film transistors and photodiodes.
Current state-of-the-art digital sensors are more sensitive
or responsive to diagnostic energy x-rays than radiographic film, and thus much less radiation (up to 70%
less) is often only required to produce a digital image.
Digital Radiography Systems Advantages and Disadvantages.
over CR. Some of these include:
• Lower doses—the PSP plates used in CR have a lower
efficiency of detection compared to DR detectors.
Thus, a higher radiation dose with CR is needed to
obtain adequate image quality.
• Greater ease of use and faster patient throughput
• Immediate imaging results: CR requires the cassette
be removed from the x-ray machine and then placed
into a reader. This is a labor-intensive step that requires the technologist to leave the patient and
workstation with each imaging procedure, even if
for a short time.
• Additional image manipulation
• Much less overall maintenance (e.g., avoidance of
CR cassette testing, cleaning, and storage space
requirements)
There are, however, several disadvantages of DR rela-
tive to CR, namely10:
• DR is much more costly
• CR is compatible with a wide range of preinstalled
traditional systems
• Multiple cassette sizes with CR allow for greater flexibility than the single detector size of DR
• The CR PSP imaging plates are subject to mechanical
damage and also to chemical oxidation but can be
replaced without great expense. DR receptors, while
well protected from mechanical damage, do not last
DR systems offer several advantages
forever; they experience gradual aging processes, and
their replacement cost is high.
An additional prospective disadvantage of DR relative
to CR is that most DR systems either do not allow the
user to change the grid to accommodate the imaging
task or have a preinstalled grid that is not easily accessible. These conditions can result in the use of grids for
pediatric imaging, thereby unnecessarily increasing a
child’s radiation dose. Because of this, DR facilities need
to collaborate with their radiation safety officer (RSO)
and physics group more than ever before to ensure the
highest-quality low-radiation dose imaging for the
smallest patients. During acceptance testing of a new DR
system, a technologist should inquire whether or not
grids are removable from the digital imaging equipment.
Repeat Rates in Digital Imaging
Because image contrast and overall light intensity may
be manipulated after image acquisition, digital imaging eliminates the need for almost all repeat images
required as a result of improper technique selection
(Fig. 11.15). However, repeat rates for reasons of poor
positioning are not lowered. Since, for DR, the image
receptor is part of the imaging equipment and does
not need to be removed for processing, the technologist can simply view the image on a monitor at the
control panel. This raises a concern regarding the
number of repeats required due to mispositioning.
There is no direct “penalty” to a technologist for
repeating images because of poor technique, so the
examination and resultant radiation exposure could
potentially be delivered multiple times, instead of just
once, without the knowledge of supervisors.
Early reports have indicated an increase in repeat rates
in DR imaging. The increase is fundamentally due to the
ease of repeating an image. It appears that technologists are
in many cases seeking to obtain the perfect image, without
the consideration of increased exposure to the patient.
Digital imaging systems, both DR and CR, separate
the process of acquisition from the display and thereby
permit a relatively large degree of viewing manipulation
of the raw images by the radiographer and/or the radiologist upon examination. Since digital devices also
typically have wide exposure latitude,* this can result in
*Exposure latitude refers to how much an image receptor can
be overexposed or underexposed and still be able to yield a
useful image.
CHAPTER 11 Equipment Design for Radiation Protection
1.6 mAs/70 kVp 3.2 mAs/70 kVp 6.4 mAs/70 kVp 12.5 mAs/70 kVp 25 mAs/70 kVp
A
211
2.5 mAs/70 kVp 5 mAs/70 kVp 10 mAs/70 kVp 40 mAs/70 kVp 80 mAs/70 kVp
B
Fig. 11.15 (A) The images obtained with a screen-film image receptor system illustrate how changing techni-
cal exposure factors greatly affects film image quality. (B) Computed radiography (CR) images obtained
through the same technique ranges as those used for (A) have much less effect on image quality because
“CR image contrast is constant, regardless of radiation exposure.” (From Betsy Shields, Presbyterian Hospital, Charlotte, North Carolina. In Bushong SC: Radiologic science for technologists: physics, biology and
protection, ed 11, St. Louis, 2017, Elsevier.)
a wide range of received patient doses, from very low to
extremely high.
An “appropriate” patient exposure is the lowest value
that is needed to provide a resultant image of sufficient
quality for a radiologist to confidently make an accurate
differential diagnosis. However, except for extreme
overexposure or underexposure, both DR and CR
images can be manipulated, because of the separation
between acquisition and display, to exhibit satisfactory
contrast resolution sensitivity. This is predominately
due to the ability of digital detector systems to adjust or
rescale the received projection data to a useful grayscale
viewing scope or range. Thus, it may be so, in many
situations, that the patient has needlessly received a
larger radiation exposure often without the knowledge
of anyone involved in the diagnostic reading of the case.
212
In some exceptional situations, an overexposure factor
of three or more might happen. In general, it has quite
often been discovered that a phenomenon known as
“dose creep” occurs due to the lack of negative impact
observed when the patient’s anatomy is overexposed but
the display still exhibits quality images.
To remedy potential overexposure and thereby adhere
to as low as reasonably achievable (ALARA) practices,
either each image taken by a digital system could be
monitored by an independent quality control technologist at a separate monitor, which might be quite tedious,
or a quality control exposure counting system could be
devised whereby for each technologist the number of
images per examination is compared with the quantity
ordered.
Alternatively, there is, however, a safety feature installed with most digital detector systems: an exposure
index (EI) indicator, derived from data collected with
anthropomorphic phantoms, that supplies the relative
speed and sensitivity of the digital receptor to incident
x-rays. The EI value ideally provides a guide to the technologist regarding the proper radiographic techniques
to select that will yield an optimal image for a specific
examination in terms of both acceptable image quality
and patient radiation dose. Usually, during commissioning of a new digital system, a table of appropriate
EIs will be determined for that system.
CHAPTER 11 Equipment Design for Radiation Protection
RADIATION SAFETY FEATURES OF FLUOROSCOPIC EQUIPMENT, DEVICES, AND ACCESSORIES
Fluoroscopic Procedures and Patient Irradiation Rates
Fluoroscopy is a continuous irradiation process that
demonstrates dynamic motion of or through selected
anatomic structures (e.g., a stomach filled with barium
sulfate and air during an upper gastrointestinal series)
by generating and displaying real-time imaging of those
structures on a monitor that works in conjunction with
an image signal amplification system. Fluoroscopic procedures (Fig. 11.16) produce the largest patient radiation exposure rates in diagnostic radiology. Therefore,
the referring physician should carefully evaluate whether
the potential benefit to the patient, in terms of information gained, outweighs any adverse somatic or genetic
Fig. 11.16 Fluoroscopic procedures produce the largest pa-
tient radiation exposure rate in diagnostic radiology.
effects of the examination. If the fluoroscopic procedure
is necessary, every reasonable effort must be taken to
minimize patient exposure time.
Fluoroscopic Imaging Systems: Non-digital
Brightness of the Fluoroscopic Image and Patient Absorbed Dose.
Non-digital fluoroscopy (Fig. 11.17) involves the use
of a signal amplification device called an image intensi-
fier (II) tube (depicted in Fig. 11.18 and discussed in
great detail in the next section). This apparatus markedly increases the brightness or intensity of the realtime image produced on a screen during fluoroscopy. II
fluoroscopy has three significant benefits, which are
listed in Box 11.4
The x-ray image intensifier (II) converts the pattern
of x-rays transmitted through the patient into a corresponding and amplified visible light pattern. With an II
the overall illumination of the raw fluoroscopic image is
increased to roughly 10,000 times that of the image obtained by non-II fluoroscopic systems* while operating
*A pre-image intensification fluoroscopic system consisted of
the fluoroscopic tube, mounted beneath the radiographic table, producing x-rays that passed through the tabletop and the
patient before striking a zinc-cadmium sulfide (ZnCdS)
screen. The latter then phosphoresced, producing an image of
the anatomy of interest. Unfortunately, this image was very
dim, and to be adequately discerned by the radiologist, a photographic dark-room situation always had to be created.
CHAPTER 11 Equipment Design for Radiation Protection
Ceiling-mounted
radiographic
X-ray tube
213
Flat panel
monitors
Image intensifier
Spot-film
cassette
Bucky slot cover
Fig. 11.17 Image intensification fluoroscopy unit. The x-ray tube used in this unit is mounted beneath the
unit’s radiographic table, which supports the patient. The image intensifier and other image detection devices
are then drawn forward and placed over the patient on the table to perform the examination. Other fluoroscopic equipment arrangements are possible. (From Bushong SC: Radiologic science for technologists: phys-
ics, biology and protection, ed 10, St. Louis, 2013, Elsevier.)
tube
Variable
aperture
collimator
Protective curtain
Fluoroscopic x-ray tube
under table
Technologist
control
Cassette tray
for overhead
radiography
’s
under the same conditions. This very large increase in
brightness has dramatically improved the radiologist’s
perception of the fluoroscopic image and at the same
time enabled a decrease in patient absorbed dose. Image
intensification fluoroscopy requires much smaller milliamperage than does pre-II fluoroscopy (approximately 1
to 1.5 mA and even less can now be used for many procedures whereas 3 to 5 mA and more was usually required before image intensification fluoroscopy). The
resultant decrease in exposure rate, for similar examina-
tion durations, yields a sizable dose reduction for the
patient.
Image Intensifier Tubes and Magnification.
An II tube is an “electronic device that receives the
image-forming x-ray beam and converts it into a visiblelight picture of high intensity.”1 A simple diagram of this
tube with components labeled is shown in Fig. 11.18.
Magnification, or multifield, features are present in the
vast majority of II systems. They are also found in digital
fluoroscopy (DF) units (see the discussion on DF pre-
sented later in this chapter). Multifield image intensification tubes vary in size, but the 30/25/20 cm (12/10/8
inch) diameter tri-field model is typical for generalpurpose fluoroscopic units. However, other sizes and
214
CHAPTER 11 Equipment Design for Radiation Protection
Output phosphor
Anode
Focal point
Electrostatic
lenses
Electrons
Glass
envelope
Fig. 11.18 Basic components of an image intensifier tube.
(From Bushong SC: Radiologic science for technologists: physics, biology and protection, ed 10, St. Louis, 2013, Elsevier.)
Photocathode
Input
phosphor
BOX 11.4 Benefits of Image
Intensification Fluoroscopy
• Increased image brightness
• Saving of time for the radiologist
• Patient dose reduction
If magnification in the fluoroscopic image is desired,
the viewing mode can be changed to the smaller 25 cm
mode (10 inches) or even less (e.g., 20 cm or 8 inches in
many new systems). With this selection, the voltage on
the electrostatic focusing lenses is increased, thereby
causing the focal point of the electrons to move closer
to the input phosphor surface or a greater distance away
from the output phosphor.1 As a result, only photoelectrons from the central 25 cm, or 20 cm, diameter portion of the input phosphor, instead of its entire surface
area, actually reach the output phosphor of the image
intensifier. This added distance from the focal point localization of the photoelectrons to the output phosphor
surface, however, creates a larger image but with a de-
creased FOV (Fig. 11.19).
In magnification mode, the quality of the enlarged
image, if there are no other changes, as viewed on
a monitor, is somewhat degraded. This reduction in
image clarity occurs because of the decrease in minifica-
tion gain (i.e., an increase in brightness) when fewer
photoelectrons are available to strike the output phosphor on the image intensifier. Therefore, the resultant
image is dimmer. Because it is necessary and desirable
to maintain a constant level of light intensity on the
25-cm
focal point
30-cm
focal point
magnification modes are also available.* When the
normal viewing mode of 30 cm (12 inches) is used,
photoelectrons emitted from the entire surface of a convex shaped cesium iodide (CsI) phosphor (i.e., the input
surface when the x-ray photons passing through the
patient first strike the II assembly) are accelerated
and converge on a focal point from which they subsequently spread out again and advance to the surface
of a zinc-cadmium sulfide output phosphor as shown in
Fig. 11.18.
*Many current II systems are available with four selectable
viewing modes or magnifications: e.g., 20 cm, 17 cm, 15 cm,
12 cm: 12”, 9”, 6”, 4” are other available selections.
25 cm
17 cm
Fig. 11.19 A 30/25/20 image intensifier tube produces a mag-
nified image in 25 cm mode, whereas the 20 cm mode produces an image that is even more highly magnified. (From
Bushong SC: Radiologic science for technologists: physics, bi-
ology and protection, ed 10, St. Louis, 2013, Elsevier.)
CHAPTER 11 Equipment Design for Radiation Protection
215
view monitor, fluoroscopic mA increases automatically
to counter this illuminance loss. The overall quality of
the image will now be enhanced relative to the largerdiameter modes because a greater number of x-ray
photons are being used to form the magnified image.
This image will have a more even appearance (less
noise), and it will be possible to distinguish among
similar tissues more readily because of improved contrast. However, the increase in tube mA necessarily
raises the dose to the patient. Thus, magnification
modes are only used when diagnostically needed.
An image intensifier can also be designed to inter-
actively change the input FOV from a large to a smaller
area. If the input FOV is halved, then the region of the
patient being observed is also halved, which results in
two-fold magnification of the image.
Pulsed Fluoroscopy. Pulsed, or intermittent, fluoros-
copy involves manual or automatic periodic activation
of the fluoroscope x-ray tube by the fluoroscopist,
rather than continuous activation. This practice:
• Significantly decreases patient dose, especially in
long procedures
• Helps extend the life of the tube
In pulsed mode the system software automatically
turns the radiation beam on and off at an operatorselected repetition rate. The most common rates are 30,
15, and 7.5 pulses per second, with each radiation pulse
lasting no more than 10 milliseconds. Shorter pulse
durations lessen the effects of any patient motion
thereby improving the sharpness of the image but decrease the signal-to-noise ratio because of the smaller
number of x-rays involved. A 30 p/s rate would typically be used for imaging studies involving very rapid
anatomic or process motion (e.g., interventional catheter procedures) to achieve acceptable temporal resolution. Barium swallow studies, on the other hand, could
make use of 7.5 p/s. Many systems include a last image
hold feature that allows the fluoroscopist to momentarily halt the radiation and review the most recent
image before giving the patient another pulse of radiation. Frequently utilizing the last image hold feature
during long procedures will noticeably reduce patient
absorbed dose.
Limiting Fluoroscopic Field Size. The radiologist must
limit the size of the fluoroscopic field to include only the
area of clinical interest. This involves visually moving
the shutters placed between the x-ray tube and the patient to define the desired field of view. When fluoroscopic field size is limited, patient integral dose (i.e.,
volumetric dose) decreases substantially. For lengthy
examinations, however, it is possible to spread out the
patient entrance irradiation area and thereby minimize
the possibility for skin effects while maintaining the
same anatomic field of view. This spreading out is accomplished by moving or rotating the patient so that
the radiation enters at multiple portions of the patient
surface instead of just the same limited region during
the procedure.
The selected primary beam length and width must
always be confined within the image receptor boundary. Regardless of the distance from the x-ray source
to the image receptor, the useful beam should not
extend outside the image receptor. Ideally, visible borders ought to appear on the image monitor. If this is
not so, a patient could receive a substantial radiation
dose in certain procedures to sensitive areas adjacent
to the study area. Thus, this conformity (alignment
and congruence) between the x-ray field and the input
phosphor of the II is an essential item of regulatory
concern.
Radiation Delivery Factors
Selection of technique exposure factors for adult patients. During manual fluoroscopic procedures, the
fluoroscopist must select technical exposure factors
that will minimize patient dose. Increasing the kVp and
filtration reduces the patient exposure and dose rate.
Most fluoroscopic examinations employ a range of 75
to 110 kVp for adult patients, depending on the body
area being examined. This kVp range produces the correct level of fluoroscopic image brightness. Lower kVp
increases patient dose because a less penetrating x-ray
beam necessitates the use of a higher milliamperage to
obtain adequate image intensity. The operator can further limit excessive entrance irradiation of the patient
by ensuring that the x-ray SSD is not less than 38 cm
(15 inches) for stationary (fixed) fluoroscopes and not
less than 30 cm (12 inches) for mobile fluoroscopes (Carms). A 30-cm (12-inch) minimal distance is required,
but an increased distance to 38 cm (15 inches) is preferred for all current systems. In addition, positioning
the II input phosphor surface as close as is practical to
216
CHAPTER 11 Equipment Design for Radiation Protection
the patient will further reduce the patient’s entrance
exposure and dose rate.
Selection of technique factors for children. Quality
radiation safety practices with fluoroscopic procedures
for children necessitates a decrease in kVp by as much as
25%. The kVp chosen should depend on anatomic part
thickness, just as it does in radiography. In addition to
decreasing kVp, maintaining SSD and minimizing the
height of the II entrance surface above the patient will
further limit excessive entrance irradiation of the pediatric patient.
Filtration. The function of a filter in fluoroscopy, as
in radiographic procedures, is to reduce the patient’s
skin absorbed dose from soft x-rays. Adequate layers
of aluminum equivalent material placed within the
collimator assembly in the path of the useful beam
remove the more harmful lower-energy photons from
the beam by absorbing them. A minimum of 2.5 mm
total aluminum equivalent filtration must be permanently installed in the path of the useful beam of the
fluoroscopic unit. With current systems, a total aluminum equivalent filtration of at least 3.0 mm or
markedly higher is typical. Although increases in filtration cause a loss of fluoroscopic image brightness,
using higher kVps can somewhat compensate. As in
radiography, the minimum permanent filtration of
the x-ray beam is federally mandated, and therefore
the HVL of the beam must be measured to confirm
agreement with regulatory standards. In routine fluoroscopy, an x-ray beam HVL of 3 to 4.5 mm aluminum is considered acceptable when kVp ranges from
80 to 100.
Cumulative Timing Device. A cumulative timer must
be provided and used with each fluoroscopic unit. This
resettable device measures the collective x-ray beam-on
time and sounds an audible alarm or, in some cases,
temporarily interrupts the radiation until it is reset after
the fluoroscope has been activated for 5 minutes. It also
serves to alert the fluoroscopist to the amount of time
the patient has been receiving x-ray exposure. Activating the fluoroscope for shorter periods will not only
cause the patient to receive less radiation exposure but
also the fluoroscopist and the radiographer. Total fluoroscopic beam-on time should be documented for every
fluoroscopic procedure.
Entrance Irradiation Rate Limitations. Current federal
standards limit entrance skin irradiation rates* of
general-purpose intensified fluoroscopic units with
maximum technique factors engaged to a maximum
of 88 mGya per minute entrance absorbed dose rate (or
10 R/min entrance exposure rate), as measured at the
tabletop using an average phantom, with the II entrance
surface at a prescribed 30 cm (12 inches) above the tabletop. This limit for conventional fluoroscopic systems has
been imposed to provide protection against patients
accidentally receiving skin-damaging entrance dose
levels in short periods. Fluoroscopic units equipped
with high-level control (HLC), however, may produce a
skin entrance dose rate as great as 176 mGya per minute
(20 R/min entrance exposure rate). Because certain
lengthy fluoroscopic procedures can result in the largest
patient doses in diagnostic x-ray imaging, sometimes
reaching the level of therapeutic doses with potential
physical damage to the patient, a concerted effort must
be made to keep fluoroscopic exposure rates and cumulative exposure times within established limits.
Primary Protective Barrier. A primary protective bar-
rier (i.e., shielding in the direct line of the transmitted
patient irradiation) of 2 mm lead equivalent is required
by regulatory guidelines for a fluoroscopic unit. The II
assembly or a digital detector is designed to provide this
barrier to direct radiation. The assembly must be physically joined with the x-ray tube and interlocked so that
the fluoroscopic x-ray tube cannot be activated when
the II or other detection system is in a parked, unaligned, and unconnected position.
*One must be precise in the terminology used for entrance or
tabletop fluoroscopy radiation levels, e.g., there is the term,
entrance or tabletop exposure rate, which has been and still
is specified by regulatory agencies and medical physicists in
R/min, and on the other hand there is the usage of mGya/min
which is correctly described as entrance or tabletop absorbed
dose rate and sometimes even as exposure dose rate. Fluoroscopic dose rate and exposure rate are not the same. There is a
numerical conversion factor existing between them, called the
rads (cGy) per roentgen factor “f,” which is about 0.88. The f
factor reflects the fact that not all incident or entrance x-ray
photons interact with and deliver energy to surface tissue atoms. Thus, an entrance exposure rate of 10 R/min in reality
corresponds to an entrance absorbed dose rate of about 8.8
cGya/min or 88 mGya/min.