Molecular imaging for visualization and quantification of ... - Philips

Investigations and research
Molecular imaging for visualization and quantification
of individualized targeted cancer therapy
M. Lubberink
S.N.F. Rizvi
O.S. Hoekstra
G.A.M.S. van Dongen
Targeted therapy and imaging
Recent advances in molecular biology have
facilitated the identification of molecular targets
on tumor cells, such as those involved in
proliferation, differentiation, apoptosis, and
angiogenesis. This has boosted the design of
targeted pharmaceuticals, with monoclonal
antibodies (MAbs) forming the largest category.
Examples of MAbs used in targeted therapy
include:
• Trastuzumab (Herceptin ® ) directed against
the HER-2 receptor and approved for treatment
of metastatic breast cancer
• Cetuximab (Erbitux ® ) and panitumumab
(Vectibix ® ) directed against the epidermal
growth factor receptor (EGFR) and approved
for treatment of e.g. colorectal cancer
• Bevacizumab (avastin ) directed against the
vascular endothelial growth factor (VEGF)
and approved for treatment of colorectal cancer
and non-small cell lung cancer
• Rituximab (Rituxan ® ) and 90 Y-ibritumomab
tiuxetan (Zevalin ® ; see Figure 1) directed
against CD20 and approved for treatment of
non-Hodgkin’s lymphoma.
These MAbs may be used as monotherapy, but
more often in combination with chemo- or
radiotherapy. Hundreds of new MAbs are under
development. The global market for MAbs is
expected to triple between 2005 and 2010, to
20-30 billion US dollars. Targeted therapies are
highly expensive and will generally be beneficial
to only a subgroup of patients, depending on
such factors as over-expression of the target,
differences in metabolism, and variability and
heterogeneity of tumor uptake. In order to
understand the efficacy of a certain targeted drug
in an individual patient, its uptake in the tumor
and normal tissues should be assessed during a
scouting procedure prior to the start of therapy or
immediately upon the start of a course of therapy.
Response to therapy is usually addressed by
anatomical (CT) imaging using RECIST criteria.
Department of Nuclear Medicine and PET Research, VU Medical Center,
Amsterdam, the Netherlands.
Departments of Nuclear Medicine and PET Research, and Otolaryngology/
Head and Neck Surgery, VU Medical Center, Amsterdam, the Netherlands.
However, anatomical changes to treatment are
usually not visible during the earliest stages of
therapy, and several cycles of therapy are given
before response can be assessed. Targeted
therapeutic agents have compromised the
association of volume-based measures and patient
outcomes even further. There is increasing
evidence that using PET with a metabolic tracer
such as 18 F-fluorodeoxyglucose (FDG) shows
changes in tumor glucose metabolism before
anatomical changes are visible. However, even
the use of FDG-PET for response monitoring
still requires one or two cycles of therapy before
the efficacy of the therapy can be evaluated.
Measurement of tumor uptake of the targeted
drug, labeled with a positron-emitting nuclide,
may provide a predictor of treatment efficacy
even before the start of therapy. A PET study
with the labeled drug could then be used to
select in advance those patients who will benefit
from a certain targeted therapy, avoiding
unnecessary administration of expensive targeted
drugs and avoiding delay in effective treatment
for the patient. This, of course, assumes that
the kinetics of a tracer amount of the drug are
similar to those of pharmacological dosages, so
that a tracer study can predict pharmacological
distribution of the tracer, and, equally important,
that the uptake of the drug in the tumor is an
accurate measure of treatment efficacy.
Quantitative imaging of targeted drugs is also of
value in drug development. Information on the
optimal dosage, the uptake in critical organs, and
interpatient variations in kinetics and targeting
can be obtained at early stages during drug
development, thereby allowing for early selection
of promising drug candidates and reducing
development costs [1].
Labeling of the targeted drug with a positronemitting
isotope allows for quantitative PET
imaging of the distribution and kinetics of the
drug both prior to and during therapy. The
radioactive half-lives of the most commonly used
PET isotopes, 18 F and 11 C, are 110 and 20.4 min,
E
E
Recent advances in
molecular biology
have facilitated the
identification of tumor
cells.
Quantitative imaging
of targeted drugs is
also of value in drug
development.
MEDICAMUNDI 54/2 2010 41

E
Table 1. Positron-emitting isotopes
for labeling of monoclonal
antibodies
1
G
Figure 1. [ 18F]FDG (left) and 6 days
p.i. [ 89Zr]ibritumomab tiuxetan
([ 89Zr]Zevalin) PET image of a patient
with Non-Hodgkin lymphoma showing
high Zevalin uptake in parailiac lymph
nodes. FDG uptake in the lymph
nodes is only moderate (SUV 2).
Images were acquired on a Gemini
TF-64 PET-CT scanner (Philips
Healthcare).
42 MEDICAMUNDI 54/2 2010
respectively. This is too short to measure the
kinetics of intact MAbs.
The longer half-life positron emitters such as
66 Ga , 64 Cu [2], 86 Y, 76 Br [3], 89 Zr [4-6] and 124 I
[7, 8] have been used for labeling of MAbs,
MAb fragments and peptides [4-6, 9-15] and a
number of mainly 124 I-labeled tracers have been
suggested for measuring apoptosis [16, 17], insulin
receptors [18], hypoxia [19], and proliferation
[20]. The half lives of these isotopes are shown
in Table 1.
Isotope Half-life
66 Ga 9.5 h
64 Cu 12.7 h
86 Y 14.7 h
76 Br 16.7 h
89 Zr 78 h
124 I 100 h
Targeted therapy with
radionuclides
Several radiopharmaceuticals have gained routine
acceptance for radionuclide therapy, with 131 I-iodide
for the treatment for thyroid cancer being the
best known example [21].
Other examples are the use of 131 I-labeled MIBG
(meta-iodobenzylguanidine) for treatment of
neuroblastoma, 90 Y, 177 Lu or 111 In-labeled
somatostatin analogues for treatment of
neuroendocrine tumors [22], and palliative or
adjuvant treatment of skeletal metastases, for
example using 89 Sr-chloride [23]. Several MAbs
labeled with 90 Y or 131 I, such as 90 Y-labeled
ibritumomab and 131 I-labeled tositumomab
(Bexxar ® ) have been approved for (radio-)
immunotherapy [24].
The biological effect of therapy with radionuclides
is due to the deposition of energy of ionizing
radiation per unit mass of tissue, i.e. the absorbed
dose. This is a well-defined physical quantity,
unlike those of other systemic treatments such
as chemotherapy [25]. Although absorbed dose
alone might not be sufficient to fully predict
response to radionuclide therapy, with dose rate,
type of radiation, and biological characteristics
of targeted drug, tumor and normal organs being
additional factors that can affect response,
knowledge of absorbed doses is required for an
optimal application of radionuclide therapy.
The amount of radioactivity injected should be
chosen such that toxicity to the dose-limiting
organ (usually red marrow in the case of
antibodies, and the kidneys in the case of smaller
molecules such as peptides) is limited, while
maximizing tumor absorbed dose.
Internal emitter dosimetry
In current clinical practice in targeted
radiotherapy, radioactivity is usually administered
as a fixed amount, or at best as a function of
patient size. Absorbed dose calculations, if done
at all, are generally performed on a total body or
region of interest basis. The necessary parameters
for internal emitter dosimetry are the number
of decays in each tissue of interest or voxel
(cumulative activity) and the geometry of the
patient which defines the transport of radiation.
Generally three to four measurements of the
radioactivity distribution can be made (Figure 2),
after which a (dual-)exponential fit through these
few data points is used to calculate the cumulative
activity in each tissue of interest, and absorbed
dose calculations are made using the MIRD
(Medical Internal Radiation Dose Committee)
approach applying a standardized geometry
[26, 27].
In more advanced methods, the geometry used
for absorbed dose calculation is based on a CT
image of the patient involved. The absorbed dose
in organ or tumor due to radiation originating
from another organ or tumor is then calculated
using voxel S-values [28], dose point kernel
methods [29] or Monte-Carlo simulations [30-34]
(Figure 3).
The main error sources in absorbed dose
calculations are the use of a standardized
geometry, the computation of the cumulative
activity, and the inaccuracy of tissue radioactivity
measurements with single photon imaging [35].

2a
2b
The distribution of most radionuclides used to
deliver the absorbed dose in targeted therapy
with radionuclides can either not be imaged at
all ( 90 Y, alpha-emitters) or can only be imaged
with limited accuracy using single photon
scintigraphy ( 177 Lu, 131 I). Here, a positronemitting
analogue of the therapeutic nuclide can
be used, if available, such as 86 Y for 90 Y [13],
124 I for 131 I, 110m In for 111 In [37], or 83 Sr for 89 Sr
[23], but otherwise an isotope with similar
chemical properties may be used. An example
of this is the use of 89 Zr as surrogate for 90 Y, which
may be preferred to 86 Y since 89 Zr has more
favorable decay characteristics for imaging and,
due to its longer half-life, covers more of the
kinetics of 90 Y [6]. The most straightforward
example here is the use of 124 I-iodide as an
analogue of 131 I-iodide in thyroid cancer imaging.
In the case of thyroid cancer therapy dosimetry,
which is a relatively straightforward example
because only a small part of the body needs to be
imaged and used for dose calculations, this has
even been done on a voxel-by-voxel basis yielding
absorbed dose images [38-41] (Figure 3).
Quantitative PET with long-lived
isotopes
Figure 4 shows simplified decay schemes of 124 I,
89 Zr and, for comparison, 18 F. In contrast with
the pure positron emitters 18 F and 11 C, the
longer-lived isotopes mentioned above all emit
gamma radiation in addition to positrons. This
gamma radiation is often emitted simultaneously
with positrons, referred to as “prompt gamma
radiation”. In the decay of 124I, for example,
about 50% of all positrons (β + in Figure 4) are
1
emitted simultaneously with a 603 keV gamma
photon (γ 1 ). This additional gamma radiation
challenges quantitative image acquisition and
image quality in a number of ways.
3
Detection of essentially true coincidences of
these prompt gamma photons with each other
or with annihilation photons introduces a bias
in the images which is not corrected for by the
standard PET corrections [42-45] (Figure 5).
This bias also results in degraded image contrast
[43, 46]. Crude correction methods for this effect
have been suggested, in the form of a uniform
background subtraction [43, 47], subtraction of
a fit to the sinogram data outside the object
[43, 48], a convolution subtraction algorithm
[45], or a point-spread function subtraction [44],
but all of these methods were developed for
conventional PET scanners and they cannot be
readily used in list-mode image reconstruction
with state-of-the-art PET-CT scanners.
The single-scatter simulation scatter correction
applied on all the latest generation PET and
PET-CT scanners [49] usually includes a scaling
F
Figure 2. Serial 89Zr-cmAb U36 PET
images of a patient with oropharyngeal
tumor (indicated by arrows),
arranged (left to right) from 1, 24,
72, and 144 h after injection. Gray
scale settings were set for each
image independently, for clarity.
Adapted from Börjesson et al. [36].
Figure 2a. Increased uptake in time
of 89Zr-cmAb U36 in tumor
(indicated by arrows).
Figure 2b. Circulating 89Zr-cmAb U36 in heart and uptake in organs.
G
Figure 3. Example of 3D 131I absorbed
dose calculations in a thyroid cancer
patient based on serial PET images
with 124I and created using the
STRATOS internal radiation
dosimetry package on an Imalytics*
workstation (Philips Research,
Aachen, Germany). Image courtesy
of Dr. Bernd Schweizer, Philips
Research, and Dr. Walter Jentzen,
University Hospital Essen.
*This product is not licensed or
intended for human diagnostic or
therapeutic use.
MEDICAMUNDI 54/2 2010 43

E
Figure 4. Simplified decay schemes
of 124I, 89Zr, and, for comparison,
the “standard” PET isotope 18F. Only radiation with abundance >1%
is shown.
5
G
Figure 5. Degrading effects in PET,
from left to right: random
coincidences, scattered radiation,
and prompt gamma coincidences
where one of the annihilation photons
is detected in coincidence with a
prompt gamma photon.
Adapted from Lubberink et al. [43].
44 MEDICAMUNDI 54/2 2010
4
to match the estimated scatter contribution to
the actual events measured just outside the body.
If this scaling includes both a multiplicative as
well as an additive factor, it implicitly performs
a crude correction for a uniform bias caused by
prompt gamma coincidences as well [50, 51].
Finally, is has been shown that the distribution
of prompt gamma coincidences matches the
distribution of random coincidences rather well
[52]. Therefore, a correction method involving
subtraction of a scaled randoms sinogram could
be an accurate correction for prompt gamma
coincidences [52], possibly incorporated into
the single scatter simulation [53].
The increased singles rate due to gamma radiation
leads to increased random coincidence rates.
This can be accurately corrected for using the
standard delayed window method, but correction
for a larger random fraction increases image
noise. One option to improve image quality may
be the use of a narrower energy window, which
reduces random coincidence rates involving
higher-energy photons, such as the 603 keV
photon emitted by 124 I (Figures 6 and 7). The
Philips Gemini TF-64 system in use in our
research department allows the energy window
to be adjusted while working in the research
mode, but this is not a capability present in the
normal, clinical mode.
The fraction of detected photons with energy
outside the scanner’s energy window increases
considerably compared to positron-only emitters.
Rejection of photons outside the energy window
does contribute to dead time, but these photons
are not counted in the singles rate. Since the dead
time correction is usually implemented as a
function of singles rate, it may become
inaccurate [42, 54].
As can be concluded from the decay schemes in
Figure 4, all of these problems affecting image
quality and quantitative accuracy occur with 124 I.
For 89 Zr, however, the 909 keV photon is not
emitted simultaneously with positrons. Therefore,
with 89 Zr, an increased random coincidence rate
can be expected, but no quantitative bias.
Hence, in terms of decay radiation, 89 Zr may be
considered the optimal PET isotope for
labeling MAbs.
Future developments
An inquiry among nuclear medicine departments
in Europe revealed a high demand for new
radionuclides, especially 124 I [55]. Methods for
large-scale production of highly pure 89 Zr and
124 I, and for facile and stable coupling of these
positron emitters to MAbs, have been developed
at VU University Medical Center, Amsterdam
[10, 56, 57]. In addition, a GMP facility has been
established for large-scale production of these
isotopes in quantities exceeding 3 GBq/day, so
that these isotopes can be supplied worldwide.
Furthermore, the availability of the PET-CT
technique is increasing rapidly. For example, in
2002 there were only two PET sites in the
Netherlands, but in 2010 there are more than 30,
almost all of which are PET-CT. Similar increases
can be seen worldwide. This allows for rapid
clinical introduction of labeled targeted drugs
for individualized therapy, and for development
of the related techniques.

6
G
Figure 6. Effect of using a narrower energy window.
Figure 6a. Noise equivalent count rates, which are a measure of signal-to-noise ratio, of the Gemini TF-64 PET/CT
with 11C using the standard 440-665 keV energy window (black), for 124I using this same window (blue) and for 124I using a narrow 440-560 keV window (red). NEC rates were normalized for positron abundance.
Figure 6b. Improvement in recovery and image contrast of 124 I using the narrower energy window.
We foresee a very important role for PET in the
development and application of targeted drugs,
as recently described by Van Dongen et al. for
Mabs [58]. However, despite clinical optimism,
it is fair to state that the efficacy of current
targeted drugs is still quite limited, with benefits
for only a proportion of patients. Moreover,
costs of these novel drugs are high, and this item
became the subject of national discussions
about the right to cancer care (e.g. trastuzumab)
in the Netherlands.
Important questions are how to improve the
efficacy of targeted therapy and how to identify
patients with the greatest chance of benefit. In
other words: when, how, and for whom should
targeted therapy be reserved?
7a 7b
Quantitative imaging of targeted drugs can also
be a valuable tool at several stages of drug G
development and application. From first-in-man Figure 7. PET images of a patient with metastatic thyroid
clinical trials with new drugs it is important to cancer at 24 h after administration of 37 MBq
learn about the ideal drug dosage for optimal
tumor targeting (e.g. saturation of receptors),
the uptake in critical normal organs to anticipate
toxicity, and the inter-patient variations in
pharmacokinetics and tumor targeting. Drug
imaging might provide this information in an
efficient and safe way, with fewer patients treated
at suboptimal dose. This approach is especially
attractive when the drug of interest is directed
against a novel tumor target that has not been
previously validated in clinical trials.
Quantitative drug imaging might also be of value
to guide optimal use of FDA-approved drugs, MEDICAMUNDI 54/2 2010 45
124I acquired
on a Gemini TF-64 PET-CT scanner (Philips Healthcare).
The narrower energy window (Figure 7b) results in a
15% improvement in image contrast in the largest
metastasis (arrow) due to the decreased image background.
Figure 7a. 440-665 keV energy window.
Figure 7b. 440-560 keV energy window.
including selection of patients with the highest
chance of benefit from such drugs.
To make this happen, software tools for optimal
imaging and improved quantification of long-lived
positron emitters are urgently needed L

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