Download Principles of X-Ray Imaging

1
Principles of X-Ray Imaging
Already a few weeks after the discovery of X-rays in 1895
by Wilhelm Conrad R€
ontgen the first medical images with
photographic plates and fluorescent screens were made. This
was the origin of projection radiography and fluoroscopy.
The greatest steps forward in X-ray diagnostic radiology
since Roentgen’s observations were the development of the
image intensifier systems and then above all the announcement of computed tomography (CT) in a clinical environment
by Hounsfield at the 1972 British Institute of Radiology
annual conference. A further important step was the introduction of digital image receptors in projection radiography
during the last years. Compared to conventional film-screen
systems these receptors allow the separate optimisation of
photon detection and image processing, resulting in significant advantages for image quality and dose.
Although today projection radiography is still the most
frequent examination with X-rays the use of computed tomography increases rapidly, and – because it involves larger
radiation doses than the conventional imaging procedures
(cf. Table 10.1) – contributes significantly to the annual
collective dose (see Fig. 1.1). Therefore CT also obtains
growing attention in radiation protection (Brenner and
Hall 2007).
In X-ray diagnostic radiology the image is generated by
the interaction of X-ray photons, which have transmitted the
patient, with a photon detector. These photons can either be
primary photons, which have passed through the tissue without interacting, or secondary photons, which result from an
interaction along their path through the patient. The secondary photons will in general be deflected from their original
direction and result in scattered radiation.
The basic principles of projection radiography/fluoroscopy
and CT are shortly explained in Sects. 1.1 and 1.2 respectively. Although totally different in image character, both
imaging systems have in common certain features, which
can be recognised in Fig. 1.2:
1. X-rays are produced in an X-ray tube.
2. The energy distribution of the photons is modified by
inherent and additional filtration.
3. The X-rays are attenuated differently by the various body
tissues.
4. Scattered radiation, which impairs image contrast, is
reduced.
5. The transmitted photons are detected.
6. The image is processed and – in the case of CT –
reconstructed.
This makes it possible to discuss the aspects of image
quality and radiation exposure for both systems together in
the main parts of the book (cf. Chap. 2).
In radiography/fluoroscopy with digital image receptors
and in computed tomography the digital image consists of a
(typically square) matrix of picture elements (pixels) which
represent the corresponding volume elements (voxels) and –
after the exposure – carry the local intensity information (gray
scale value). Quality of digital images depends primarily on
the image matrix size, i.e. the pixel size (cf. Chap. 9). As the
matrix size is increased resolution improves but the number
of photons in each pixel must be increased in order to maintain
a certain minimum noise level.
1.1
Projection Radiography
and Fluoroscopy
In projection radiography and fluoroscopy the image is a
two-dimensional projection of the attenuating properties of
all the tissues along the paths of the X-rays. The components
of a typical radiographic/fluoroscopic system are shown in
Fig. 1.3.
The photons emitted by the X-ray tube are collimated by
a beam-limiting device. Then they enter the patient, where
they may be scattered, absorbed or transmitted without
interaction. The primary photons recorded by the image
receptor form the image. The secondary photons create a
certain amount of background radiation which degrades
contrast. If necessary, the majority of the scattered photons
can be removed by placing an anti-scatter device between
the patient and the image receptor. This device can simply be
H. Aichinger et al., Radiation Exposure and Image Quality in X-Ray Diagnostic Radiology,
DOI 10.1007/978-3-642-11241-6_1, # Springer-Verlag Berlin Heidelberg 2012
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1 Principles of X-Ray Imaging
Angiography and
intervention
2%
Mammography
4%
CT
7%
Remainder
0.7%
Dental
0.2%
Remainder 1% 3% Thorax
9%
Skeleton
GI and urogen
and bile tract
8%
Dental
3%
GI and urogen
and bile tract
37%
1%
Mammography
CT
60%
Skeleton
33%
18%
Angiography
and
intervention
Thorax
13%
Fig. 1.1 Contribution of various examination types to total frequency (left) and to collective effective dose (right) in 2006 for Germany adapted
from BMU (2009)
X-ray tube
x
Collimation
Production of X-rays
Filtration
Object transmission
Patient
Scatter reduction
Photon detection
Image reconstruction and processing
Fig. 1.2 Basic principles of radiography/fluoroscopy and CT imaging
Patient support
Anti-scatter device
AEC system
Image receptor
Fig. 1.3 Typical arrangement of a radiography/fluoroscopy system
1.2
Computed Tomography
an air gap or a so-called anti-scatter grid formed from a
series of parallel metal strips. An automatic exposure control
system (AEC) provides for the correct exposure of the image
receptor. Today digital image receptors predominate in radiography and fluoroscopy, but film-screen systems and image
intensifiers are also still in use.
1.2
Computed Tomography
Whereas it is not possible in projection radiography to
gain any depth information from a single image, computed
tomography separates the superimposed anatomical details
and produces sectional or axial slice images with excellent
soft tissue contrast. Compared to projection radiography and
fluoroscopy computed tomography is a rather new imaging
technique. Therefore it seems to be reasonable to present its
fundamental principles in some more detail.
The principle of computed tomography is illustrated in
Fig. 1.4.
A well-collimated X-ray pencil beam is attenuated by
the tissues along its path and the transmitted radiation is
detected. In order to generate one projection the tube-detector
assembly scans the object in a linear translatory motion. This
procedure is repeated at many viewing angles (typically at
least 180 projections are received with a rotational increment
of 1 ). From these projections a two-dimensional discrete
distribution of the linear attenuation coefficients mtissue is
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reconstructed as image signal by computation. In practice
CT numbers or Hounsfield units are used instead of mtissue
where the Hounsfield unit HU is defined by:
HU ¼ 1000 ðmtissue mwater Þ
mwater
(1.1)
where mwater is the linear attenuation coefficient of water.
The experimental set-up of Hounsfield corresponded
largely to the arrangement sketched in Fig. 1.4.
This set-up was termed the ‘first generation’ of CT
(Kalender 2006). To speed up scanning and to utilise the
available X-ray power more efficiently the first commercial
scanners (the ‘second generation’) used some more detectors
and a small fan beam. The typical scan time for an 80 80
image matrix was 5 min (Kalender 2006).
Continuously rotating CT systems (‘third generation’)
according to Fig. 1.5 with a fan beam covering the total patient
cross-section and a corresponding detector array, consisting
of gas proportional detectors or scintillation detectors (cf.
Sect. 8.2), were introduced in the 1980s. Continuous rotation
was made possible by a slip-ring technology for electrical
power supply and data acquisition. Scan time was reduced
down to 2 s for a single slice with a 256 256 matrix.
A major step forward in CT technology was the introduction of spiral or helical CT by Kalender and Vock in 1989
(Kalender et al. 1989; Vock et al. 1989): Slice-by-slice
imaging was replaced by volume scanning. The principle
X-ray tube
Detector
Fig. 1.4 Principle of data acquisition in CT imaging (Adapted from
Bunke 2003)
Fig. 1.5 Continuously rotating CT system with a fan beam and
corresponding detector array (Adapted from Bunke 2003)
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1 Principles of X-Ray Imaging
Fig. 1.6 Principle of spiral CT
imaging (From Bunke 2003)
of this method is illustrated in Fig. 1.6: While the fan beam is
continuously rotating the patient is moved with constant
velocity along his body axis (the z-axis) through the gantry;
this results in a spiral track of the focal spot around the
patient and accordingly in a spiral data set.
Direct image reconstruction from these data would give
rise to image artefacts (similar to motion artefacts). This
can mostly be avoided by data interpolation. The interpolation
method developed at first was the 360 linear interpolation
(LI) algorithm, which used data from a full rotation of the
tube-detector assembly. Since for a complete interpolated data
set at a definite slice position two successive 360 rotations on
either side of the selected plane were necessary, considerable
widening of the slice profile resulted, thus reducing image
quality. Therefore the 360 LI was soon replaced by a 180
LI where interpolation from opposing 180 points reduces the
spiral range used for reconstruction. This is possible since
X-ray beam attenuation at a distinct rotation angle j is equivalent to the X-ray beam attenuation traversing the body from
the opposite side, at 180 + j. As the distance of the data
points is now smaller, effective slice width will be less.
In 1992 CT scanners were introduced, which used two
parallel banks of detectors. This was followed by multirow
detector CT scanning in 1998 using solid detectors and simultaneously imaging four slices in each rotation of the X-ray tube
(Kalender 2006). A great advantage of multislice CT (MSCT)
scanners over single section spiral CT is the opportunity for
longer anatomic coverage during the same scanning time.
[ mm ]
5
2.5
1.5 1 1 1.5
2.5
5
2×8
4×5
4 × 2.5
4×1
2 × 0.5
Fig. 1.7 Adaptive array detector with detector combinations for
different slice thicknesses (from Bunke 2003), e.g. the uppermost
combination allows slice widths in the longitudinal direction from
1 to 5 mm at the isocentre
The MSCT detector arrays could be divided into two
groups: Those with detector elements of unequal width along
the z-axis (adaptive array detector) and those with elements of
equal width (linear or matrix detector). Figure 1.7 shows as an
example an adaptive array detector with the possibility of the
setting of different slice thicknesses.
At present 64-slice scanning represents the state of the art,
allowing the imaging of all body regions with submillimetre isotropic spatial resolution and scan times of 5–15 s
(Kalender 2006).
References
Scan time can be further reduced with recent developments
of CT such as dual source CT or cone beam CT with C-arm
systems (Kalender 2006). This is especially interesting for
cardiac imaging, angiography and interventions. Dual source
CT scanners are equipped with an ultrafast dual detector
system and two X-ray tube assemblies. Cone beam CT (see
Sect. 8.2) uses a flat-panel-detector with up to 1,920 rows
and 2,480 columns (Oppelt 2005). This enables enhanced use
of X-ray quanta, but also leads to a higher fraction of scattered
radiation (see Sect. 11.2.4).
References
Brenner DJ, Hall EJ (2007) Computed tomography – an increasing
source of radiation exposure. N Engl J Med 357:2277–2284
Bunke J (2003) Computertomographie. In: Schmidt T (ed)
Strahlenphysik Strahlenbiologie Strahlenschutz. Springer, Berlin,
pp 84–98
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Kalender WA (2006) X-ray computed tomography. Phys Med Biol 51:
R29–R43
Kalender WA, Seissler W, Vock P (1989) Single-breath-hold spiral
volumetric CT by continuous patient translation and scanner rotation. Radiology 173:414
BMU (Bundesministerium f€
ur Umwelt, Naturschutz und Reaktorsicherheit) (2009) Umweltradioaktiuit€at und Strahlenbelastung im
Jahr 2008: Unterrichtung durch die Bundesregierung http://nbnresolving.de/urn:nbn:de:0221-201003311019
Oppelt A (ed) (2005) Imaging systems for medical diagnostics.
Publicis, Erlangen
Vock P, Jung H, Kalender WA (1989) Single breathhold spiral volumetric CT of the lung. Radiology 173:400