Download medical x-ray imaging, current status and some future

Copyright ©JCPDS-International Centre for Diffraction Data 2006 ISSN 1097-0002
MEDICAL X-RAY IMAGING, CURRENT STATUS
AND SOME FUTURE CHALLENGES
Erik L. Ritman, M.D., Ph.D.
Department of Physiology & Biomedical Engineering
Mayo Clinic College of Medicine, 200 First Street SW, Rochester, MN 55905
ABSTRACT
Biomedical x-ray imaging has proven to be an indispensable component of many medical
diagnostic and treatment techniques. Major technical advances over the years, culminating in
helical scanning, multi slice, x-ray computed tomography, can now provide ∼ 0.5 mm3 resolution
3D images of the spatial distribution of tissue attenuation coefficients of an entire adult human
thorax or abdomen. The utility of the x-ray images is ultimately limited by two opposing factors,
one is that increased image spatial and contrast resolution requires an increased number of
detected x-ray photons per image resolution unit, and the other is the direct relationship between
the number of x-ray photons and the increased risk of tissue damage and ultimate cancer
development. This trade-off limits the utility of x-ray imaging in screening of subjects who are
likely to be at risk of a disease (e.g., a family history or exposure to an environment that
increases the risk of a life-threatening disease) but who have not yet developed the clinical signs
or symptoms. Based on increasing experience (especially with synchrotron x-ray-based
imaging), x-ray imaging of some consequences of a tissue’s x-ray refractive index, rather than
x-ray attenuation, has potential for decreasing the x-ray dose and increasing contrast resolution.
This approach, while presenting formidable technological problems for whole-body 3D imaging,
could, if achievable in a routine clinical setting, greatly expand the applicability of x-ray imaging
for screening and repeated studies of asymptomatic subjects of all ages.
INTRODUCTION
X-ray has been used for biomedical imaging ever since Roentgen x-rayed his wife’s hand in
1895 [1]. The speed with which the technology was implemented in medicine is phenomenal [2]
by today’s standards. This is well illustrated by a note written by Edison on January 27, 1896 to
a Mr. Kennelly, which said “How would you like to come over and experiment an Rotgons (sic)
new radiations. I have glassblower and pumps running and all photographic apparatus. We
could do a lot before others get their second wind” [3]. An important reason for this rapid
acceptance is probably that it was the first time that objective evaluation of an internal organ
could be made in time to act in a meaningful way, and the other reason is probably that the
technology for making x-rays (i.e., high voltage electricity and the Crooke’s tube) and recording
the x-ray images (i.e., photographic plate) were available because they were already in wide
spread use for other purposes. Nonetheless, resistance to x-ray imaging soon followed its
introduction to medical practice [4] and continues in recent times as part of effort to control
spiraling medical costs by controlling the introduction of high technology into medical practice
[5].
1
This document was presented at the Denver X-ray
Conference (DXC) on Applications of X-ray Analysis.
Sponsored by the International Centre for Diffraction Data (ICDD).
This document is provided by ICDD in cooperation with
the authors and presenters of the DXC for the express
purpose of educating the scientific community.
All copyrights for the document are retained by ICDD.
Usage is restricted for the purposes of education and
scientific research.
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Copyright ©JCPDS-International Centre for Diffraction Data 2006 ISSN 1097-0002
The history of the development of x-ray imaging has been explored in a number of excellent
books [6], so only a brief overview of the developments in the past century is given here. Since
Roentgen, the methodology for making x-ray images has improved considerably. The
introduction of Coolidge’s rotating x-ray anode tube [7] allowed for great increase in x-ray flux
so that the duration of the exposure was diminished to milliseconds which eliminated motionblurring, and also increased the number of patients who could be imaged in a day. Similarly, the
use of anti-scatter grids and increased x-ray film sensitivity and resolution also contributed [8].
The introduction of x-ray contrast agents that selectively highlighted the contents of
physiological spaces such as the blood vessels, bowel lumen, cerebrospinal fluid space etc., also
increased the spectrum of abnormalities that could be detected [9]. The developments of the
television camera tube, image intensifier and videotape recording with instant stop-action replay
capability, have enabled the explosion of x-ray-guided interventions such as treatment of
localized arterial narrowing by balloon dilatation and stenting.
Computed tomography [10] has greatly increased the specificity and sensitivity of image
information content [11] because this approach has eliminated superposition and foreshortening
of anatomic structures. In the past 15 years fast, multi-slice, helical scanning CT [12,13]
(Figure 1) has rapidly advanced clinical CT imaging.
Figure 1 – Schematic of helical
CT scanning. This involves
continuous rotation of the x-ray
source and its detector array
about the patient while the table
is continuously advanced.
With the introduction of multi-slice (up to 64) helical CT scanning, the field of preventive
screening has been opened up. Examples include virtual colonoscopy [14], lung imaging for
early detection of cancer [15] and coronary artery calcification for detection of presymptomatic
coronary atherosclerosis [16], as illustrated in Figure 2.
However, while there is still considerable potential for further technical improvement of the
helical CT methodology , especially in terms of increased speed, spatial resolution and volume
scanned, the approach seems to be fast approaching the limits of its clinical “utility” envelope, if
not its “technical potential” envelope. The reason is basically that today’s x-ray methodologies
can involve near-therapeutic levels of x-ray exposure [17,18]. As a consequence it is unlikely
that the contrast discrimination, spatial resolution, repetition rate of x-ray images and volume of
the body imaged can be improved significantly using refinements of current technologies
involved in x-ray imaging.
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Figure 2 – Left panel: Computer generated display of 3D image generated with a multi-slice helical CT
scanner during intra venous injection of iodinated contrast medium. The image is useful for identifying
narrowed or blocked coronary arteries. The middle panel shows cross-sections of a coronary artery, the
bright lumen (due to contrast medium in lumen) and gray wall surrounding it. The right panel also shows a
very bright region of calcification within the arterial wall. The ability to image the arterial wall has great
potential for early detection of arterial disease before symptoms arise. [Left panel: courtesy of Dr. Thomas
C. Gerber, Mayo Clinic Jacksonville, FL. Copyright 2005, Mayo Foundation for Medical Education and
Research. Right panel: E.L. Ritman: Cardiol Clinic 21:491-513, 2003].
To overcome this “barrier” to progress we need to look at both the ultimate biomedical needs and
requirements as well as the technical possibilities, such as:
1) What radiation doses are acceptable?
2) What spatial, contrast and temporal resolution are needed?
3) What volume of the body needs to be imaged concurrently?
4) What 3D image repetition rate is needed?
5) Which, more specific and effective, contrast media need to be developed?
6) How much can prior information contribute?
7) How much can technological improvements be expected to contribute?
8) Are there x-ray/matter interactions, other than attenuation, that could be used to meet clinical
needs at orders of magnitude lower radiation dose?
X-RAY DOSE CONSIDERATIONS
Radiation exposure should be limited to keep deleterious biological effects to a minimum [18].
However, for a fixed signal to noise ratio in each CT image voxel (3D pixel), the x-ray exposure
needs to increase by at least the cube of the resolution, i.e., by increasing from 1 to 0.5 mm3
resolution the exposure should increase at least 8 fold [19]. Just 0.1 Grey [10 Rad] absorbed
dose in tissue is likely to ultimately result in cancer in up to 0.1% of people exposed [20]. In the
average patient 2 Grey absorbed dose will result in erythema, (reddening of the skin) and 18
Grey will result in tissue necrosis and ulceration over several weeks. To put x-ray exposures in
context, a single scan sequence of the thorax, using a multi-slice helical CT scanner, involves an
exposure of 0.03-0.07 mSv/second when scanning at the level of the mid abdomen [60] which
results in absorbed dose rates to the skin that are typically 2-8 mGy/second. Annual exposure at
sea level from cosmic and terrestrial sources is an average, 3.6 mSieverts and a chest x-ray is
0.04 mSieverts [21].
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SPATIAL AND CONTRAST RESOLUTION
Up until recently the basic diagnostic marker in x-ray images was change from “normal” and/or
change from a prior x-ray image. With the advent of CT the absolute gray-scale of the image has
increasingly played an important role beyond just the discrimination of different tissues [22].
Current CT images can discriminate gray and white matter in the brain and abnormally high
levels of iron accumulation in the liver. However, the contrast range in conventional CT of
different soft tissues (e.g., muscle, blood, brain) is much more limited than that in Magnetic
Resonance Imaging (MRI).
Of particular interest is the use of altered image gray-scale due to introduction of contrast
medium (usually iodinated sugar-like molecules) both as an index of abnormal accumulation in a
physiological space where it would normally not enter (e.g., indicating a leaky blood vessel such
as occurs in cancerous tissue) and as a means of performing indicator dilution analysis for
characterizing transport phenomena such as blood flow [23]. Up until recently most contrast
agents injected into the blood stream were selectively excreted by the kidneys. Fortunately, this
phenomenon also provides an index of renal function (and anatomy) as well [24].
Figure 3 illustrates selective opacification of the liver by use of recently developed 200 nm
diameter chylomicron-like particles [25]. This is an example of a contrast agent that is
selectively accumulated and excreted by the bile-producing cells in the liver.
Figure 3 – Cryostatic micro-CT image of a
rat liver obtained two hours after
intravenous injection of poly-iodinated
triglyceride nano-particles Fenestra®. The
bright areas are hepatic lobules in which
the hepatocytes have incorporated the
contrast agent for excretion into the bile.
Spatial and contrast resolution are inversely inter-related and often conveyed in terms of the
Modulation Transfer Function [26]. Factors that affect spatial resolution include x-ray source
focal spot size and shape, detector thickness and “granularity” (i.e., actual microstructure of the
scintillating x-ray-to-light converter) and CT image-generation factors such as number of angles
of view, reconstruction algorithm characteristics etc. Contrast resolution [27,28] depends on
signal noise (especially in x-ray quantum-limited images) and contamination by scattered x-ray
which adds noise and a variable bias component [29]. Another contributor to altered gray scale
values is beam-hardening due to a polychromatic (bremstrahlung) x-ray beam becoming
“hardened” (i.e., less attenuating) as it traverses tissue, and especially regions containing material
such as bone that have increased attenuation [30,31]. Increasing the x-ray exposure reduces the
fraction of the measured (i.e., transmitted) signal that is noise and a goal is to have the
modulation of transmitted signal be five times greater than the noise for unequivocal detection of
the signal [32]. Use of a large object-to-detector distance can greatly reduce the scatter that is
detected, but with finite focal spot dimensions this is at the cost of reduced spatial resolution [33]
due primarily to the penumbral blurring effect. Use of mechanical collimators preferentially
reduces the x-ray scatter contamination of the image and thereby increases the contrast resolution, but at considerable cost of loss of signal by obscuration by the collimating structure [34].
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The Basic Functional Unit (BFU) of organs seems like a reasonable goal for the spatial
resolution clinical whole-body imagers. BFUs are the smallest accumulations of diverse cells
that function like the organ they are in, e.g., the hepatic lobule in the liver or nephron in the
kidney. Generally, they occupy about 0.01 mm3, i.e., about 200 μm in diameter, which is
essentially true for all organs and for all vertebrate animals. The BFU volume is essentially the
upper limit of a spherical piece of tissue which can survive without its own blood supply.
Hence, early ‘solid’ cancers would grow to this size before a blood supply is needed to sustain
further growth [35]. Thus, an isotropic spatial resolution of 100 μm seems like a reasonable goal
because it would allow us to measure their number, size and perhaps individual functional status.
However, because they, and their microvascular blood supply, consist of tissue that is
elementally not very different from the connective tissue in which the BFUs are embedded, this
might involve very high radiation exposures to provide the signal to noise ratio needed for
discriminating a BFU from its surrounding tissue matrix. Because their contrast differential is
very low, either suitable contrast agents that selectively accumulate in (or avoid) them, will need
to be developed (e.g., such as illustrated in Figure 3), and/or an x-ray imaging method which has
inherently greater contrast (than does conventional, attenuation-based, x-ray imaging) will need
to be developed.
VOLUME TO BE IMAGED
The volume to be scanned is determined by the need to image an entire organ. This is because
the absence of a detected abnormality in an organ will only then have meaning. This would be
the case for a tumor (due to cancer or some other focal pathology) or some local breakdown of
the organ that could initiate a disseminating cascade of tissue failure [36,37]. The gut would
require that the entire abdomen is imaged and the lung requires that the entire thorax is imaged –
both being approximately 30-40 cm in cephalo-caudal height and in transverse diameter.
TEMPORAL RESOLUTION
In addition to 3D anatomy, the rate, amplitude, synchrony, and the spatial distribution of the
movement attributes of organs are important indicators of the organs’ function(s). There are two
types of involuntary motion; one can be predictively cyclic, such as the heart beat or breathing,
and the other is unpredictable, and generally not reproducible, such as the movement of the
bowel or the passage of a bolus of contrast agent through the blood stream. Thus, there are two
issues; one is to image fast enough to stop the motion blurring i.e., the 3D imaging “aperture”
time, which would need to be fast enough to stop-action the fastest moving phase of the motion;
the other is the 3D image repetition (or frame) rate needed to quantitate the motion. The heart
wall normally has a peak rate of thickening of approximately 50 mm/sec so that a stop-action
exposure duration would need to be at most ([spatial resolution]/ [50 mm/sec]) seconds which is
about 10 msec at 0.5 mm resolution, although if the imaging is to occur only during the slowest
motion phase of the cardiac cycle (end diastole) that would be about 100 msec. In the case of the
predictable cyclic motion, such as occurs in the heart, incremental sets of partial scan data can be
acquired progressively over many sequential cycles by acquiring the scan data during a fixed
phase of that cyclic motion. This “gated” scanning allows imaging of, for instance, the entire
heart at one phase of multiple sequential cardiac cycles (i.e., over a scan period of one breath
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hold). The other issue is the number of sequential exposures needed to adequately describe the
motion. For a cardiac cycle, which in an adult is approximately half to one second in duration,
for a breath this might be 5 seconds, for a blood-borne indicator dilution curve this might be 20 –
30 seconds and for a bowel motion study this might be 30-60 seconds. However, the fastest
motion within that cycle would determine the image “frame” rate needed – for the left ventricle
of the heart this about 15 images/second [38]. However, detector lag resulting in mixing of
current and previous images can also greatly degrade the image quality if the image moves
during the sequential exposures.
X-RAY/MATTER INTERACTION OPTIONS EXTENDING X-RAY IMAGING
INFORMATION
With current multi-slice helical CT scanners the x-ray volume scanned at approximately 0.5 mm3
voxel resolution can encompass the entire abdomen or thorax within a single, ∼ 20 second,
breath-hold and, if ECG-gated scanning is used, 170 millisecond temporal resolution can be
achieved. Thus, the main limitations of that approach are that the spatial and temporal
resolutions are still not optimal and that the x-ray exposure is approaching unacceptable levels
for repeated screening purposes in asymptomatic subjects. Past experience, using a
one-of-a-kind CT scanner, has shown that 20 cm diameter cylindrical volumes can be scanned in
11 milliseconds and repeated at up to 60/second [39] and experience with high flux synchrotron
radiation has shown that micrometer resolution dynamic x-ray projections can be generated [40].
These approaches have been useful for research applications but have not evolved into clinically
useful scanners. So the question is, can the desired resolution characteristics be achieved by
further technical developments of the attenuation-based imaging approach?
1) OPTIMIZATION OF ATTENUATION-BASED IMAGING BY USE OF:
a) Monochromatic x-rays instead of polychromatic (bremstrahlung) x-rays [41,42] would limit
the radiation exposure to just the useful x-ray photons and also make the image contrast
information more quantitative. In favorable circumstances this may reduce x-ray exposure to
half. If in addition photon energy discrimination could be used, most scattered photons could be
identified by their lower photon energy and thus rejected without the need for mechanical
collimation (thereby eliminating that source of reduced x-ray detection efficiency).
b) K-edge absorption discontinuities subtraction imaging to selectively enhance contrast has
been shown to be very effective [43] in terms of enhancing signal to noise. However this
approach has been limited by the difficulty of generating the desired monochromatic x-rays at
the intensity and photon energy needed for routine clinical use. The effectiveness of contrast
agents can be greatly enhanced if the absorption K-edge of the element in the contrast agent is
tightly “straddled” by the dual energy x-ray exposures. While very effective, as has been
demonstrated with iodine-based contrast agent, the presence of contrast in large physiological
spaces (e.g., cardiac chambers – 5 cm diameter) overlying the structure of interest (e.g., coronary
artery lumen – 4 mm diameter) can greatly reduce signal to noise [44]. Contrast agents-based on
elements such as Europium, would involve ∼47 keV photon energies that better penetrate
through the body than does the ∼ 31 keV needed for iodine-based contrast agent.
c) Counting individual photons [45]. This would eliminate the “noise” involved in the current
method of integrating the signal detected in the x-ray-to-light/current converter (with an analog
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circuit). If, in addition, photon energy discrimination could be used, then many scattered
photons could be detected and rejected without the need for mechanical collimation.
d) Scattered x-ray to provide added image information because the mode of scatter contains
information not conveyed by the attenuation. However, this wouldn’t reduce radiation so much
as provide more information per unit of x-ray exposure [46]. As a large proportion of x-ray
photons in clinical x-ray imaging is scattered, it might be possible, in selected applications, to
obtain the desired information by use of scatter detection involving lower x-ray exposures than
would be needed for attenuation-based images of the same structure or its dynamics [47].
e) Inversion of the single x-ray focal spot/imaging array approach to an array of x-ray sources
illuminating a single detector approach has the advantage that the entry x-ray exposure, in terms
of photons/mm2 of skin is reduced, even though the total radiation dose to the body remains
unchanged [48].
f) Hardware and software ‘tricks’ that can be used to reduce patient exposure. One method is to
use “bow tie” shaped attenuator which spatially modulates the x-ray exposure so that the
transmitted x-ray signal is more uniform in its spatial distribution, thereby increasing the
effective dynamic range in the image. Another involves varying the radiation exposure at
different angles of view so as to keep the transmitted x-ray signal constant at the level needed for
the desired signal to noise ratio [49]. For a chest image this means that the Lateral exposure
would be less than the Antero-Posterior exposure.
Another approach is to restrict the x-ray exposure to just the organ of interest (e.g., the heart).
This would reduce the amount of tissue exposed but needs special “limited data set” tomographic
reconstruction algorithms to provide useable images [50,51].
g) A priori information to help deduce quantitative information about the sub-resolution
structure of tissues by use of a suitable “inverse” model. Such a model would have a priori
information about the tissue microstructure, such as has been attempted for myocardial
microvasculature [52]. This approach would reduce the need for high resolution imaging in
selected application and thereby reduce the x-ray exposure needed.
2) USE OF PAIR-PRODUCTION FOR IMAGING
If x-rays with photon energy greater than 1.022 MeV are used to irradiate a body, pair production
can result and these can be detected via the positron annihilation which results in two 511 keV
gamma-ray photons being produced in essentially opposite directions. If the body is exposed in
a plane, then the detection of ‘coincident’ gamma-rays on opposite sides of that irradiated plane
directly provides the tomographic image information [53]. This method would have contrast
discrimination by virtue of the pair production being sensitive to atomic number and density of
material. Hence, it should be particularly suitable for bone imaging and with suitable contrast
agents. When the exact ‘slice’ to be imaged is known – then this approach exposes only that
slice with the illuminating beam. Because the attenuation of the 511 keV photons by the body on
either side of the imaged slice is relatively low (∼0.1/cm), the tomographic image is only
minimally degraded by attenuation of the 511 keV photons.
3) IMAGING OF X-RAY REFRACTIVE INDEX
The refractive index for clinically relevant x-rays through tissue is very close to 1, but its
variation amongst tissues theoretically provides greater contrast than does attenuation, especially
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of low atomic weight elements [54]. Moreover, the amount of x-ray exposure needed to obtain a
useable image should also be greatly reduced relative to attenuation-based imaging used to detect
the same contrast differences [55,56].
a) Refraction of the x-rays, especially at sites of very rapid change in refractive index (such as
occurs at the edges of tissue fibers or small ducts and blood vessel walls) results in refraction of
the x-ray so that at some distance beyond the refraction site an edge enhancement can occur [57].
This approach is probably limited to several centimeters of tissue depths and tissues in which
such fibrous/vascular micro-structure is an important clinical cue (e.g., mammography in which
the breast is compressed down to a 4 cm thickness). A variant of this method involves use of
grating interferometers placed adjacent to the detector side of the person (i.e., the phase grating)
and the other just in front of the detector (i.e., the absorption grating) [58].
b) Differences in phase shifts that occur due to passage through materials of different refractive
index can be used to cause Moiré patterns when ‘mixed’ with a reference beam. The Moiré
pattern can be unraveled to provide a map of the phase shift distribution [59]. This works well
for a small thickness of tissue in which a small number (e.g., 5) of phase shift occurs (and hence
can be accounted for) but after passage through 30 cm of tissue the amount of phase shift,
combined with the shorter photon wavelength needed to provide adequate penetration of the
body, would make “unraveling” of the phase shift very more difficult.
c) The ultimate goal would be to directly measure the spatial distribution of the transit time of
individual photons. As x-ray photons travel at the speed of light each photon traverses 1 mm in
approximately 3.3 x10-12 seconds, i.e., 3.3 picoseconds. However, as the velocity of x-ray
through water differs little from that through air, then imaging of just the presence or absence of
tissue at 1 mm resolution would require time discriminations in the low femto-second ranges and
that is probably not practical at the clinical level at present. Differences between tissue types
would affect the velocity of x-ray predominantly via density differences, and because
proportionately fewer of the high energy photons are attenuated, the x-ray dose received by the
subject is reduced. However, the overwhelming advantage of velocity measurement would be
that only one photon would provide this measurement (although uncertainty in the velocity
measurement would require several photons to be measured to provide an acceptable degree of
certainty) whereas attenuation estimation needs more than 104 photons to be transmitted to
provide a useful measurement per picture element (pixel) in the detector.
CONCLUSION
Current, attenuation-based x-ray imaging methods cannot be expected to increase much more in
resolution due to the undesirable x-ray exposure consequences. If time-of-flight x-ray imaging
can be developed for clinical applications, the power of x-ray imaging could be greatly enhanced
without the added radiation consequences.
AKNOWLEDGEMENTS
Dr. Ritman’s research is supported in part by NIH grants HL72255 and EB000305. Dr. Ritman
thanks Ms. Mara Lukenda for help with editing and formatting the manuscript.
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