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through the radial or femoral artery, which is then conducted to the coronary artery
where a contrast agent is injected followed by x-ray imaging (see figure 2 in [55]).
Electron beam computed tomography (EBCT), having a fast acquisition time
(100 ms/slice) and a low dose (0.7 mSv), is used to non-invasively image coronary
arteries. EBCT is generally used in assessing the coronary artery calcification scores
and volumes with electrocardiographic gating, where it was observed to reduce the
variability in the images and to monitor atherosclerosis progression and the effects
of medications.
9.4.2.3 Magnetic resonance
Magnetic resonance imaging (MRI) allows the imaging of both plaque volume and
composition. It has the potential to characterize the plaque through different
imaging techniques (T1, T2 and proton-density weighting, perfusion, diffusion
and biochemical contrasts) providing their composition: fibrous cap, lipids and
calcifications [56–58]. MRI can be used to assess the effect of a treatment by
measuring the extent of the plaque and its constituents (see figure 5 in [59]). The
interest in using conventional gadolinium-based contrast agents is useful for fibrous
cap delimitation, plaque vascularization and inflammation which could contribute
to detect plaque vulnerability. Magnetic resonance angiography allows generating
images of flowing blood in the arteries, similar to x-ray angiography, providing
information on stenosis. With 1.5 T MRI scanners, the spatial resolution in image
slices approximates 0.25 × 0.25 mm
2
in the carotid and 0.46 × 0.46 mm2in the
coronary artery [60]. With 3 T scanners, the signal-to-noise ratio and
contrast-to-noise ratio jump to 223% and 255%, respectively, and the images are
acquired at a spatial resolution of 0.31 × 0.31 × 3mm
3
in the carotid wall [61]. There
is a compromise between signal-to-noise ratio, acquisition time and spatial resolution. The time usually needed to image a carotid artery is about 45 min [62].
9.4.2.4 Optical coherence tomography
OCT is an intravascular imaging modality, based on infrared light emission at
wavelengths of 1280–1350 nm. It is like ultrasound imaging but uses light instead of
sound and the reflected light on the tissue borders is analyzed with interferometers. It
is a very fast imaging method that can be processed in the time or frequency
domains. The light is delivered through an optical fiber and the beam is directed by a
microprism. The images are obtained at 20 frames/sec at axial and lateral spatial
resolutions of 10–15 μm and 94 μm, respectively, with a maximal depth in tissue of
3 mm (see figure 8 in [63]). OCT has some limitations. For example, it cannot provide
measurements of arteries in the presence of blood; blood flow has to be stopped by
inflating a proximal occlusion balloon. OCT provides better arterial wall images
than IVUS and can detect early construction of the plaques, while it does not allow
enough penetration to visualize thick plaques.
9.4.2.5 Radionuclide imaging
Radionuclide imaging has the advantage of providing information on a given
physiological phenomenon in diverse ways. However, because it is based on
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radiation emission that could be risky to the patient, small amounts of the
radiotracer are injected into the patient, and consequently, the scanners are made
of larger detectors to count more radiations. At the end, the images present reduced
contrast and spatial resolution. The two main devices used in tomographic radionuclide imaging are PET and single photon emission computed tomography
(SPECT). PET relies on the detection of two coincident gamma rays of the same
energy, while SPECT detects single photons of different energies depending on the
isotopes used for imaging. PET imaging coupled to CT imaging will be covered in
the next section. We provide an overview in the rest of this section of some studies
performed with SPECT using different radioisotopes and radiotracers.
LDL is recognized to be involved in the formation and vulnerability of the
plaque, thus it was labeled with the isotopes
111
(t
131
In
I
= 2.8 day, Eγ = 245 keV) [65],
1/2
(t
= 8 day, Eγ = 364 keV) [66–68].
1/2
99m
(t
123
99m
Tc
I
Tc
= 6h,Eγ = 141 keV) [47, 64],
1/2
(t
= 13.2 h, Eγ = 159 keV) and
1/2
-LDL was found to be located in
lesions rich in macrophages, however, it was found not to be suitable in coronary
arteries as its specific uptake was very low and its washout was found to be very
slow, making the blood activity very high and resulting in a poor lesion-tobackground ratio [47]. Radiotracers with iodinated LDL were found to be inefficient
as they were found to rapidly deiodinate in the tissues [ 66].
Several studies reported imaging of atherosclerosis in animal models using
different radiotracers, among them
111
-LDL, which was evaluated in the aorta
In
of hypercholesterolemic rabbits and showed an uptake 2.5 times higher than in
normal rabbits [65]. Oxidized LDL was shown to be present in atherosclerotic
arteries but not in normal arteries [69, 70]. Endothelial cells, smooth muscle cells and
monocyte macrophages have been reported to oxidize LDL [71]. Labeled antibodies
against oxidized LDL with
imaging (
125
-MDA2, a monoclonal antibody against malondialdehyde-lysine
I
epitopes) [69]. The authors concluded that
125
(t
I
= 59 day, Eγ = 35 keV) was used in rabbit
1/2
125
-MDA2 has excellent uptake and
I
specificity for atherosclerosis lesions. However, its application in humans should
take into account the low energy of the isotope and other parameters related to
imaging.
Several other techniques reported imaging in atherosclerosis based on markers,
we refer here to some of them [72–84].
9.5 Imaging of atherosclerosis with PET/CT
CT imaging is not only coupled to PET to seek compensation for the low spatial
resolution of PET, but also as a complement for better assessment of atherosclerosis,
such as the grading of the calcification. Of course, images acquired with CT and
those acquired with PET in the same patient and in the same session have different
sizes. Either axially or transaxially, CT images have smaller voxels than PET
images. In order to match the two sets of images for image fusion or registration, an
interpolation is mandatory. This interpolation generally introduces uncertainties in
voxel values. Since the atheromatous plaques and the calcifications have small sizes
in comparison to the PET spatial resolution, these uncertainties drastically affect the
artery images. Most PET scanners are now equipped as bimodality PET/CT
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scanners. Note that several works on atherosclerosis have been reported using dual
modality PET and MRI [48, 85–87].
Imaging is a priori intended to visualize structures such as stenosis, wall thickness
and plaque volume, and consequently to be able to measure their extent. However,
PET imaging, in addition to its higher sensitivity to the picomolar [88], can also
report measurements of metabolism and other physiological parameters. In fact, it is
not the size of the plaque which is of importance, but its inflammation, which is of
crucial importance as an indication of its vulnerability. The degree of inflammation
can be assessed by the metabolism of the cells within the plaque. During the
measurement with PET, the acquisition can be reconstructed in a sequence of images
as a function of time. This dynamic imaging is very useful as it can help in
discriminating tissue activity from blood activity and two or more tissue types.
Moreover, within a single tissue, dynamic imaging can help in determining the state
of the radiotracer as compartments which add more precision in the quantitative
analysis. This type of measurement is more appropriate in research than in the clinic
as the patient is injected with the radiotracer when lying within the scanner and the
measurement takes a longer time than in the clinic. In the clinic, the patient is
injected with the radiotracer and remains for some time before being imaged for a
short time. In this way more patients can be imaged in the clinic.
9.5.1 Fast quantitative assessment
Several studies on atherosclerosis have been reported using PET. However, most of
these studies have been carried out retrospectively. Patients referred to PET imaging
by their practitioners for a given disease such as cancer or cardiovascular problems
generate considerable data to be exploited for other analyses. These data are
generally composed of static images, thus the only way to quantitatively extract
useful information from these data is by using simplified models. The two most
commonly used models are the standardized uptake value (SUV) and the target-tobackground ratio (TBR).
Based on the images, SUV is either calculated from a region of interest (ROI) or
from image voxels. Its value, which is unitless, depends on image intensity
injected activity
where
in Bq kg−1is the mean intensity in an ROI or in a voxel and is measured
C
PET
and patient weightw:
C
SUV
PET
=
Aw
[Bq kg ]
[Bq]/ [kg]
1
−
,
in a given interval of time in a steady state with minimal variation. Activity
is the injected activity of the radiotracer, and
w
in kg is the body weight of the
, the
C
PET
(9.1)
in Bq
subject. SUV is a simple and rapid quantitative technique, sometimes it is called
semi-quantitative. By using the same PET measurement protocol in two or more
measurements, SUV can be used, for example, to compare inflammation progress or
the effect of treatments. SUV is very helpful for cancer diagnoses. However, it is not
robust and this is a function of several factors: (1) the radiotracer is not equally
distributed in the body in tissues compared to fat and hence the division by the
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weight is not certain; (2) the time of image acquisition affects the radiotracer uptake
in the tissues which directly affects SUV; (3) the ROI or the voxel grouping are very
subjective due to uncertainty introduced by the operator and the ROI can be
affected by partial volume effect and spillover; and (4)
is a measure of the whole
C
PET
radiotracer without distinction of blood and tissue in the ROI and thus SUV cannot
discriminate between perfused and necrosed tumors or inflamed tissues. Other
factors in fl uencing SUV can be found in [89–92].
To correct for the uncertainty on body weight by taking the fat into consideration, it is usual to replace the subject’s weight by the subject’s body surface area
(BSA) given by [90]
2 0.425 0.725
=× ×wBSA[m ] [kg] height [cm] 0.007184.
(9.2)
Another form of BSA is given by [93]
21/2
=×wBSA[m ] [ [kg] height[cm]/3600] .
(9.3)
Let us see the difference between equations (9.2) and (9.3). We calculated BSA1
according to equation (9.2) and BSA2 according to equation (9.3) for the values of
the weights
w
and the heights
:w= [10 20 30 40 50 60 70 80 90 100 110 120 130];
= [100 160 170 175 180 160 185 165 160 178 182 158 173]. The plot of BSA2 as a
function of BSA1 is shown in figure 9.1 together with the regression line which has a
slope of 1.09 and an intercept of −0.15. Also they were correlated (99%) and not
significantly different (p = 0.95 at 5%). The two definitions of BSA could then be
considered to be equivalent.
TBR was used in several publications to account for radiotracer uptake in artery
tissue in comparison to an ROI drawn on a venous blood image, thus TBR accounts
for artery plaque SUV divided by venous blood SUV [94, 95]:
=TBR SUV /SUV .
Figure 9.1. Comparison of BSA1 and BSA2 as defined in equations (9.2) and (9.3).
artery vein
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(9.4)
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9.5.2 Kinetic modeling
In research, the subjects, either animals or humans, are selected based on predetermined criteria, and the imaging is performed according to a predefined protocol.
In order to extract maximal information from the images, the subjects are directly
measured at the start of the radiotracer injection. The data are then collected in list
mode which can be reconstructed later in any time framing. This dynamic
acquisition can be used with SUV and TBR by considering the optimal interval
of time of the images. Moreover, uptake of the radiotracer in an artery can be
evaluated as a curve instead of a single data point. Any image voxel is made of the
contribution of the radiotracer in blood and tissue, and this is more justified in the
artery images, adding the fact that artery tissue forms a thin layer in the image even
on CT images, and it cannot be discriminated. The mixture of blood and tissue
responses is mainly due to the limited spatial resolution of the scanner (about 3 mm)
and to the subject’s organ movement. There are, however, mathematical tools to
allow decomposing a sequence of dynamic images into two components in order to
isolate uptake of the radiotracer in artery tissue. Because the dynamic behavior of
the radiotracer concentration in tissue and in blood is different as a function of time
(two curves of different shapes), the mathematics can thus split the intensity of a
voxel into two fractions of intensities representing the two components. This
operation can be achieved by algorithms such as factor analysis of dynamic
structures (FADS) [96] and independent component analysis (ICA) [97]. The
extraction of the blood component from PET images is crucial in kinetic modeling.
Instead of sampling blood from the patient, which subsequently needs centrifugation
to extract blood plasma, measurement of the radioactivity in the samples and
calibration and timing with the PET images, the mathematically obtained image of
the blood component can be used in a straightforward manner.
We briefly describe two approaches of kinetic modeling which are more accurate
then SUV. More details can be found in [98]. The first one is compartmental
modeling and here it is expressed based on 18F-fluorodeoxyglucose (18F-FDG), an
analogue of glucose, which is limited at the phosphorylation step [99]. 18F-FDG
provides high contrast images in tissues metabolizing glucose. The second approach
is an approximation of the first one, assuming no dephosphorylation of 18F-FDG.
This approach is the graphical analysis method or, as it is commonly called, the
Patlak method [100].
The 18F-FDG compartmental model is shown in figure 9.2 from which a set of
differential equations are deduced.
These two equations can be directly used in an algorithm to compute the rate
dC t
f
dt
()
=−+ +
KC t k k C t k C t
() ( ) () ()
bfm123 4
()
dC t
m
dt
kC t kC t
() ().
=−
fm34
9-12
(9.5)
4
g
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Figure 9.2. Three-compartment 18F-FDG model.
metabolized 18F-FDG.
exchanged between the compartments by means of the rate constants
and has units in ml g
constants
−Kk
1
represents what the scanner measures in a voxel or in an ROI. The radiotracer is
C
PET
−1
min−1while
−kk
24
, or they can be analytically solved to isolate each compartment
,
and
C
b
have units in min−1.
C
C
f
are the compartments for blood, free and
m
. Note that
−Kk
14
is the perfusion
K
1
and used to output the four rate constants.
The measured radioactivity
the three compartments: blood (
). In terms of equations,
(
C
m
K
Ct
1
αα
−
21
+− +−− ⊗ +
kk e kke Ct Ct() [( ) ( ) ] () (),
341 234
Ct()
C
Ct()
PET
αα υ=
where1and2are combinations of the rate constants [99];
of the radiotracer in the plasma.
is the sum of the radioactivity emitted from
PET
), free 18F-FDG (
b
) and metabolized 18F-FDG
C
f
is expressed as
αα−−
tt
12
can be determined from blood samples or
Ct()
p
pbPET
is the concentration
Ct()
p
(9.6)
simply from the blood component image obtained from the decomposed images
with FADS or ICA or with other algorithms. The fourth parameter
accounts for
the fraction of blood measured in tissue with PET. The symbol ⊗ represents the
operation of convolution.
Generally the algorithm of Levenberg–Marquardt is used to adjust the model in
the right member of (9.6) to the PET data in the left member [101].
From the calculated rate constants, the metabolic rates for glucose (MRG) are
deduced:
(mg of glucose/100 ml of plasma)
MRG( moles/100 g/min)
μ=
gl
−
LC 0.182(mg moles )
×μ
11
−− −
(ml g min ) (min )
Kk
1
×
−−
(min ) (min )
kk
2
×
1
3
+
3
1
1
,
1
(9.7)
where
is the concentration of glucose in plasma, and LC is the lumped constant.
l
LC could be considered as unity to obtain relative values of MRG. The factor 0.182 in
the denominator accounts for the molecular mass of glucose which is 182 g mole
In some tissues such as the heart and tumors, there is no dephosphorylation back
to 18F-FDG which means there is no
graphical analysis model [98, 100]:
Ct
()
PET 1 3
CtKkkk
()
p
=
23
+
in the model. The model then reduces to the
4
t
Cudu
()
p
∫
0
Ct
()
p
Kk
+
kk
23
12
+
.
9-13
−1
(9.8)
.
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In this case, equation (9.8) appears as a linear regression with slope
13
. This
+Kkkk
23
is the same ratio as in equation (9.7). The slope is evaluated on the last few data
points where the plot is linear.
The complexity of kinetic modeling resides in the determination of the plasma
function (or as it is usually called, the input function). The other difficulty is the
uncertainty on the region to consider for kinetic modeling. Since the artery plaque is
small and contains a few voxels, the time–activity curve (
Ct()
) is noisy and,
PET
consequently, the values of MRG are obtained with high variations, which affects
the difference between the normal and inflamed artery.
9.5.3 Multiple approaches in atherosclerosis quantitation with PET
Until now, the most commonly used radiotracer to assess the vulnerability of plaque
is 18F-FDG, because the inflammation, including the aggregated cells in the plaque,
metabolize glucose. It is expected that the more the plaque uses glucose, the more it
is vulnerable. Another reason for the use of glucose is its production and availability
in every PET center. Wu et al reported that 18F-FDG uptake is a function of the
availability of MMP in the plaque, which is an indicator of advanced atherosclerosis
and plaque rupture [102]. However, although PET-18F-FDG has the potential to
measure the progress of the disease, it was shown to be in discrepancy with CT.
Uptake of 18F-FDG could be present with the presence or absence of calcification,
as seen on CT images, and the opposite is also true, i.e. the presence of calcification
on CT images with no uptake of 18F-FDG [103]. Davies et al combined PET-18FFDG with MRI in a dynamic PET acquisition of 120 min. However, the authors
calculated the ratio of radiotracer uptake in the lesions with respect to uptake in a
normal artery and they established a classification of inflammation depending on the
deviation from the mean values [85]. Instead of taking advantage of the dynamic
measurements and computing MRG values, they only used the last 30 min of the
acquisition where the radioactivity in the blood is expected to be reduced and they
averaged voxel intensity in ROIs based on MRI images.
It has been shown that 18F-FDG uptake in the arteries correlates with some
cardiovascular risk factors. Therapies with either medication or lifestyle could
decrease 18F-FDG uptake, but it has not been demonstrated to be translated to
lowering the impact of the disease or even coupled to systemic biomarkers [88, 104,
105]. Repeated measurements with 18F-FDG showed up to 50% variability between
6 months and 26 months, and it was reported that 18F-FDG uptake could remain
constant for a period of 6 months [88]. Atherosclerosis can be diagnosed in young
individuals, and it is recognized that healthy elderly people are also affected by
atherosclerosis, and the disease may progress within a year [106]. In this study, the
authors demonstrated different mean 18F-FDG SUVs in three groups of patients:
healthy, with hypercholesterolemia and with chest angina, all of them above
65 years of age. The hypercholesterolemia subjects were prescribed Rosuvastatin
and the group with angina were taking their own medication, and the measurements
were repeated 12 months later [106](figure 9.3). Their results showed on average a
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Figure 9.3. CT (left) and PET-18F-FDG (right) images of calcified aorta (arrows) in transaxial (top) and
coronal (bottom) slices.
high 18F-FDG uptake in groups 2 and 3 at first scans with respect to group 1, and
the increase of 18F-FDG uptake 12 months later was less in group 3 than in the
other two groups.
Instead of estimating glucose metabolism which a priori was not specific enough
to prevent plaque rupture, some researchers based their works on the presence of
some proteins or other substrates indicating plaque vulnerability. They investigated,
among other attempts, in particular in animal imaging, the mechanisms by which
the plaque becomes vulnerable. Several radiotracers have been utilized, among them
18F-NaF [107–109], 11C-PK11195 [110], 11C-choline, 68Ga-[1,4,7,10-tetraa-zacyclododecane-N,N′,N″,N‴-tetraacetic acid]-
D-Phe1, Tyr3-octreotate (DOTATATE)
[88, 111] and 18F-VCAM-1 [112]. Joshi et al reported that 18F-NaF binds to regions
with necrosis, macrophages, apoptosis and micro-calcification [109]. 18F-NaF is
recognized to bind to bones and produces excellent images without uptake in most
tissues. Then, what is the difference between imaging with 18F-NaF and CT as both
of them can show the calcifications? Considering their advantages, CT provides high
resolution images while 18F-NaF produces a low dose to the patient, but this is not
crucial. The response to this question was given by Joshi et al where they specified
that CT provides images of calcified arteries at the final stage of calcification, while
18F-NaF can detect the slowly developing and metabolically active calcifications
that are not apparent on CT images and in some cases not on 18F-FDG images (see
figure 1 in [109]). Other works reported the specificity of 18F-NaF to detect active
plaques (see figure 2 in [108]). They noticed large calcified structures on CT images
but with low uptake of 18F-NaF and, in other regions, they observed intense uptake
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of 18F-NaF with minimal or no CT calcification. However, the potential of
18F-NaF to prevent plaque rupture remains to be confirmed.
9.6 Discussion
Imaging technologies have evolved signficantly despite their high cost and the need
to form multidisciplinary research or clinical teams. The dual modality imaging
devices, PET/CT, PET/MRI and SPECT/CT, combine inputs from the departments
of radiology and nuclear medicine, if the latter includes PET imaging. The
specifically missing contribution is not on the hardware side, it is more methodological, where some techniques need to be implemented and analyses need to be
carried on quantitatively. Most of the studies relied on SUV for their statistics. The
other crucially missing contribution is that related to the production of radiotracers.
It is in fact costly in resources and time to produce and validate a radiotracer.
From the PET imaging side, it is possible to label a molecule to measure the
concentration, perfusion, metabolism or receptors, but the focus should be on the
atheromatous plaque if the mechanisms are understood. This means that we know
what is happening in the plaque, and we want to measure it through imaging (the
bottom-up approach). Alternatively, we do not know what exactly is happening
there, and we are trying to understand the phenomenon by imaging (a top-down
approach). Researchers are yet at the exploration stage, trying to design an
appropriate radiotracer to potentially provide accurate measurements of disease
progression.
Following these ideas, several researchers are intensively developing radiotracers
targeting LDL, antibodies, MMPs, nanoparticles, monocytes and other cells. At
present, the variability in quantitative imaging is large [86], which is more drastic in
radionuclide imaging. If a molecule specific to the vulnerable plaque is identified and
labeled with an isotope, the low spatial resolution of PET will not be a hindrance.
The imaging tools are not used for diagnoses and follow-up of atherosclerosis, they
are mostly used for research to understand the development of the disease and to
prevent cardiovascular complications.
9.7 Conclusions
The techniques reported to date are largely based on simple approaches of analyses
such as SUV and uptake of the radiotracer. This is partly because of the high
variability in such small regions of the developing plaque, even when imaged with
multimodalities. To overcome the low spatial resolution of PET, specific radiotracers to prevent plaque rupture should be designed. It is not expected that the
plaque vulnerability will be assessed as present or absent, but it is expected that this
vulnerability will be determined in grades.
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