Biomed Imaging Interv J 2007; 3(2):e28
© 2007 Biomedical Imaging and
Internal dosimetry as a tool for radiation protection of the patient in nuclear medicine
MG Stabin*,1, PhD, CHP,
1 Vanderbilt University, Nashville, Tennessee, United
2 Department of Physics, Royal Marsden Hospital, London, United
A number of radioactive therapeutic agents are currently
employed against various forms of cancer and other diseases. In radiotherapy
with external sources of radiation (including brachytherapy),
patient-individualised dose calculations are always performed prior to therapy
and form an essential basis of the patient treatment plan.
Patient-individualised dose calculations, however, are usually not calculated
to optimise this process in the therapeutic use of radiopharmaceuticals. This
article reviews the current status of dose calculations for
radiopharmaceuticals and discusses the need for patient-individualised dose
calculations to optimise therapy for patients and provide improved clinical
Standardised methods and models for internal dose
Reliable estimates of radiation-absorbed dose from the use
of diagnostic or therapeutic radiopharmaceuticals in nuclear medicine are
essential to the evaluation of the risks and benefits of their use. To estimate
absorbed dose for all significant tissues, one must determine the quantity for
each tissue. The absorbed dose is defined as the amount of energy from ionising
radiation that is absorbed per unit mass of any material. These authors�
interest in the evaluation of dose from radiopharmaceuticals is in the energy
deposited in human tissues.
Standard Dose equations
A generic equation for the absorbed dose rate in an object
uniformly contaminated with radioactivity (for example an organ or tissue with
radiopharmaceutical uptake) may be shown as:
���� =������� absorbed
dose rate to a target region of interest (Gy/sec or rad/hr)
(MBq or mCi) in source region S
� =�� number
of radiations with energy Ei emitted per nuclear transition
� =�� energy
per radiation for the ith radiation (MeV)
�� =�� fraction
of energy emitted in a source region that is absorbed in a target region
of the target region (kg or g)
��� =�� proportionality
constant (Gy-kg/MBq-sec-MeV or rad-g/mCi-hr-MeV)
The proportionality constant k includes the various factors
that are needed to obtain the dose rate in the desired units, from the units
employed for the other variables, and it is essential that this factor is
properly calculated and applied. For example, if the dose rate is wanted in
rad/hr, and there are employed units of mCi
for activity, MeV for energy and g for mass, the conversions that are needed
If instead the dose rate is wanted in Gy/sec, and there are
employed units of MBq for activity, MeV for energy and kg for mass, the
conversions needed are:
Often investigators require an estimate of total absorbed
dose, rather than just the instantaneous dose rate at some point in time,
from a radiopharmaceutical administration. In the dose equation, the quantity
activity (nuclear transitions per unit time) causes the outcome of the equation
to have time dependence. To calculate the cumulative dose, the time integral of
the dose equation must be calculated. In most cases, the only term which
depends on time is activity, so the only factor that has to be integrated is
the activity term. The integral of the time-activity curve (i.e. the area under
that curve, regardless of its shape), is often called the cumulated activity
(often given the symbol �), and it represents the total number of
disintegrations that have occurred over time in a source region.
Therefore, the equation for cumulative dose becomes:
where D is the absorbed dose (Gy or rad) and the quantity �S
represents the integral of AS(t), the time-dependent activity
within the source region:
where A0 is the activity administered to
the patient at time t = 0, and fS(t) may be called the
fractional distribution function for a source region (fraction of administered
activity present within the source region at time t). In many instances, the
function fS(t) may be modeled as a sum of exponential functions:
where terms f1�fN� represent
the fractional uptake of the administered activity within the 1st to
Nth compartments of the source region, l1�lN
represent the biological elimination constants for these same compartments, and
lP represents the
physical decay constant for the radionuclide of interest. Other functional
expressions may be used to represent the fractional distribution function, but
exponentials are most commonly encountered.
A generalised expression for calculating internal dose,
which may describe the equations shown in publications by different authors,
can be calculated by the following equation:
where N is the number of nuclear transitions that occur in
source region S (identical to �S), and DF is a �dose factor�. The
factor DF contains the various components shown in the formulas for S and SEE
SEE (and in some presentations may as well include a �radiation weighting factor�, wR);
it depends on combining decay data with absorbed fractions (AFs), which
are derived generally using Monte Carlo simulation of radiation transport in
models of the body and its internal structures (organs, tumours, etc.):
As written, the above equations give only the dose from one
source region to one target region, but they can be generalised easily to
multiple source regions:
Available body models
The current generation of anthropomorphic phantoms began
with the development of the Fisher-Snyder phantom , which employed a
combination of geometric shapes - spheres, cylinders, cones, etc. - to create a
reasonably accurate representation of the body. Monte Carlo computer programs
were used to simulate the creation and transport of photons through these
various structures in the body, whose atomic compositions and densities were
based on data provided by the International Commission on Radiological
Protection (ICRP) in its widely quoted report on �Reference Man� , now
updated in a more recent report . These reports provide various anatomical
data helpful in producing dose calculations for standardised individuals.
Absorbed fractions and dose conversion factors (S values), as defined above,
for over 100 radionuclides and over 20 source and target regions, were also
published [4, 5].
Cristy and Eckerman  modified the adult male model and
developed models for a series of individuals of different size and age. Six
phantoms were developed, which were assumed to represent children of ages 0
(newborn), 1, 5, 10, and 15, and adults of both genders. Absorbed fractions for
photons at discrete energies were published for these phantoms, which contained
approximately 25 source and target regions. Tables of S values were never
published, but ultimately were made available in the computer software called �MIRDOSE� ,
which was widely used by the nuclear medicine community. Stabin et al.
developed a series of phantoms for the adult female, including a model of the
non-pregnant adult female, and the woman at three stages of pregnancy. These
phantoms modeled the changes to the uterus, intestines, bladder, and other
organs that occur during pregnancy, and included specific models for the fetus,
fetal soft tissue, fetal skeleton, and placenta. S values for these phantoms were
also made available to the dosimetry community through the MIRDOSE software.
Marrow dose models
Spiers et al. at the University of Leeds  first developed
electron absorbed fractions (AFs) for bone and marrow for an adult male
subject; these results were used to calculate dose conversion factors (DCFs),
or S values, in MIRD Pamphlet No. 11 . Eckerman  re-evaluated this work
and extended the results to derive dose factors for 15 skeletal regions in six
models representing individuals of various ages. The results were also used in
the MIRDOSE 3 software  to provide average and regional marrow dose, and
dose-volume histograms for individuals of different ages. Bouchet et al. 
used newer information on regional bone and marrow mass, and calculated new AFs
using the EGS4 Monte Carlo code. Although the results of the Eckerman and
Bouchet et al. models were similar in most characteristics and reported
results, the models differed in a few important underlying assumptions. A
revised model, which resolves these model differences in ways best supported by
currently available data, has been derived . New skeletal average absorbed
fractions for all bone regions employed in the calculations in this study were
implemented in the OLINDA/EXM  computer code, designed as a successor to
the MIRDOSE code .
Planar methods for quantification
In the introduction to MIRD Pamphlet No. 16 , the
following is stated:
�To determine the activity-time profile of the radioactivity
in source regions, four questions need to be answered:
What regions are source regions?
How fast does the radioactivity accumulate in these source regions?
How long does the activity remain in the source regions?
How much activity is in the source regions?
The first question concerns identification of the source
regions while the second and third questions relate to the appropriate number
of measurements to be made in the source regions as well as the timing of these
measurements. The fourth question is addressed through quantitative external
counting and/or sampling of tissues and excreta.
Each source region must be identified and its uptake and
retention of activity as a function of time must be determined. This provides
the data required to calculate cumulated activity or residence time in all
source regions. Each region exhibiting significant radionuclide uptake should
be evaluated directly where possible. The remainder of the body (total body
minus the source regions) must usually be considered as a potential source as
well. Mathematical models that describe the kinetic processes of a particular
agent may be used to predict its behavior in regions where direct measurements
are not possible, but where sufficient independent knowledge about the
physiology of the region is available to specify its interrelationship with the
regions or tissues whose uptake and retention can be measured directly. The
statistical foundation of a data acquisition protocol designed for dosimetry
requires an adequate number of data points and careful selection of the timing
of these points. As the number of measurements increases, the confidence in the
fit to the data and in the estimates of unknown parameters in the model is
improved. As a heuristic or general rule of thumb, at least as many data points
as the number of initially unknown variables in the mathematical curve-fitting
function(s) or in the compartmental model applied to the data set, should be
obtained. For example, each exponential term in a multiexponential
curve-fitting function requires two data points to be adequately characterised.
On the other hand, if it is known a priori that the activity retention in a
region can be accurately represented by a monoexponential function,
restrictions on sampling times are less stringent as long as enough data points
are obtained to derive the fitted function. Because of problems inherent in the
collection of patient data (e.g., patient motion, loss of specimen, etc), the
collection of data above the necessary minimum is advisable.�
The execution of a successful dosimetry study lies in the
gathering of sufficient data to characterise the radiopharmaceutical kinetics,
and in the use of those image data to identify the important source regions and
assign activity levels to them. To determine radiation dose, the counts seen in
the images must be converted to absolute values of activity (Bq or mCi), which
requires a known calibration factor for the camera, and collected data permit
correction of the raw images for radiation attenuation and scatter. In planar imaging,
the external conjugate view counting pair (anterior/posterior) as the most
common method used to obtain quantitative data for dosimetry. In this method,
the source activity Aj is given by the expression:
where IA and IP are the observed
counts in the anterior and posterior projections (counts/time), t is the
overall patient thickness, μe is the effective linear
attenuation coefficient, C is system calibration factor C (count
rate per unit activity), and the factor f represents a correction for the
source region attenuation coefficient (μj) and source thickness
(tj) (i.e., source self-attenuation correction). This
expression assumes that the views are perfectly collimated (i.e. they are
oriented towards each other without offset) under the model of narrow beam
geometry without significant scattered radiation effects. Corrections for scatter
are usually necessary, and a number of methods have been proposed. One
relatively straightforward correction procedure for scatter compensation
involves establishing adjacent windows on either side of the photopeak window,
with the area of the two similar adjacent windows equal to that of the
photopeak. The corrected (true) photopeak counts CT are given by the
where Cpp is the total count recorded within the
photopeak window, while CLS and CUS are the counts within
the lower and upper scatter windows, respectively. If the areas of the scatter
windows are not equal (in sum) to that of the photopeak window, then an
appropriate scaling factor (FS) should be applied. Subtraction of
the adjacent windows is assumed to compensate for the high-energy photon
scatter tail upon which the true photopeak events are superimposed. Even if the
areas of the scatter windows are equal to that of the photopeak window, use of
a scaling factor other than unity may provide the best correction for scatter
in a given system with a particular radionuclide. This may be determined by the
study of a known volume source in a water phantom whose dimensions are similar
to that of a human subject.
Other corrections are often required as well. Whenever a ROI
is drawn over a source region on a projection image, some counts from the
region will contain counts from activity in the subject�s body that is outside
of the identified source, scattered radiation from other ROIs, background
radiation, and other sources. A background ROI is usually drawn over some
region of the body that is close to the source ROI and which, in the
investigator�s judgment, best represents the underlying and overlying tissues
in which the source resides and which will provide the best estimate of a
background count rate to be subtracted from the source ROI. Background ROIs
should not be drawn over a major blood vessel or other body structure that
contains a high level of activity, as this will remove too much background from
the source ROI. The exact prescribing of locations and sizes of background ROIs
is very difficult, and methods vary considerably between investigators. This
can lead to markedly different results for the final estimates of activity
assigned to a source ROI. This process should be carried out with caution and
attention given to the above points for the best and most reproducible results.
For quantification of counts in the total body, or in the check source placed
externally to the body, a ROI should be drawn away from the subject�s body,
also away from any �star pattern� streaks that may accompany the source image,
but close enough to capture a typical number of counts per pixel that
represents background and scattered radiation within the imaging area close to the
It is not uncommon for some organs or tumours to have
overlapping regions on projection images. The right kidney and liver are
frequently partially superimposed on such images, as are the left kidney and
spleen, for example. When organ overlap occurs, an estimate of the total
activity within a source can be obtained by a number of approximate methods.
For paired organs, such as kidneys and lungs, one approach is to simply
quantify the activity in one of the organs for which there is no overlap with
other organs, and multiply the number of counts in this organ by two to obtain
the total counts in both organs. Another approach is to draw a ROI over the
organ region in scans where there is overlap, count the number of pixels and
note the average count rate per pixel, then use a ROI from another image in
which there is no apparent overlap and the whole organ is clearly visible;
count the number of pixels in a larger ROI drawn on this image, and then
multiply the count rate per pixel from the first image by the number of pixels
in the second image. Or, equivalently, take the total number of counts in the
first image and multiply by the ratio of the number of pixels in the second to
the first image ROIs. If a significant overlap of images with another organ is
not possible, an approximate ROI may need to be drawn just from the knowledge
of the typical shapes of such organs. This kind of approximation is obviously
not ideal, but may be necessary.
In addition, calibration coefficients for each radionuclide
and gamma camera/collimator combination must be obtained by imaging a small
source of known activity for a fixed amount of time. The attenuation
characteristics of the camera may be studied by imaging this source with
various known thicknesses of tissue-equivalent material interposed between the
source and camera, and fitting the results (counts versus thickness) to an
Quantification of tomographic data
Tomographic imaging offers the potential for improved
dosimetric accuracy due to its increased contrast when compared with planar
imaging. Tomographic data are particularly useful for dosimetry where there is
suspected heterogeneous uptake of activity in the source organ or underlying or
overlying background activity. To date, Positron Emission Tomography (PET) data
have been little used for dosimetry, although PET quantification is an active
area of research in its own right. Standardised uptake values (SUVs) are used
to quantify radiotracer uptake (FDG) and, whilst prone to some uncertainty, are
nevertheless used clinically with more regularity than quantification of Single
Photon Emission Computed Tomography (SPECT) or planar data. SUVs are defined as
Quantification of SPECT data is accomplished in a number of
Data acquisition: Data acquisition parameters include the number
and timing of scans, as detailed above. With allowance made for adequate
statistics and spatial resolution, no special criteria are required for the
number of projections or matrix size, which are usually according to standard
local protocols. However, it is worth noting that for many therapy scans,
patients will have a relatively high level of activity so that camera
sensitivity is not an issue. Data acquired soon after administration for
example, may require scan times of only 5 seconds per view, so a full scan can
be acquired in less than 10 minutes, which will not impact greatly on the daily
routine of the nuclear medicine department. Collimators should be carefully
chosen to suit the radionuclide being imaged.
Deadtime corrections: Scintillation cameras are designed for use
with low levels of Tc-99m and so are poorly adapted for use with high
activities of, for example, I-131. A common problem with these systems is that
of deadtime, whereby the counts registered do not increase linearly with the
activity in the field of view. Both paralyzable and non-paralyzable systems may
be characterised for their deadtime behavior so that correction factors that
can be directly applied to the image data may be obtained.
Scatter correction: Scatter correction methods are generally
performed by the application of scatter windows placed adjacent to the
photopeak. Frequently triple energy window (TEW) scatter correction is
performed, as detailed above, although dual energy windows are also employed. A
number of authors have explored the use of a large number of scatter windows
 or the acquisition of list mode data 
although limitations are often imposed by the system used.
Attenuation correction: Attenuation correction is a crucial step
in the quantification of emission data and a number of techniques have been
used. The most basic method employed is that of Chang et al  who assume
that the imaged volume is uniformly filled with water. More complex solutions,
that to date have seldom been used in clinical practice, involve adaptation of
a patient CT scan to provide an attenuation map .
Reconstruction: Image reconstruction can be divided into filtered
back projection and iterative techniques, and can incorporate scatter and
attenuation correction. Each of these techniques has a number of variables that
may be employed, including smoothing parameters and the number of iterations.
The effect on quantification of adjusting these parameters should be studied
carefully on a case-specific basis.
Quantification: Conversion of counts to absolute values of
activity may be performed in a number of ways. As specified above for the
processing of planar data, it is possible to include a source of known activity
within the field of view at the time of scanning. An approach used by some
authors is to construct calibration phantoms that are aimed to emulate the
patient. However, the ideal approach is to characterise the camera system so
that conversion factors may be obtained without recourse to phantoms or
external sources while taking into account patient CT data.
Quantification of image data has been considered for many
years, although as yet there are no standardised methods for quantifying SPECT
or PET data. This remains the largest single obstacle to accurate dosimetry,
and is currently a strong focus of research [16, 19, 20] . It is probable that
this task will be made easier with the advent of dual modality scanners and it
is hoped that in time manufacturers will develop systems that are adapted to
high energy high activity imaging, whereby camera sensitivity may be sacrificed
to some extent in favour of spatial and energy resolution.
Absorbed Dose and the Biologically Effective Dose (BED)
The intermittent and sparse application of dosimetry for
targeted radionuclide therapy (TRT), and the wide variation of methods
employed, mean that to date few correlations between calculated absorbed doses
and either therapeutic response or degree of toxicity have been seen. A further
complication is caused by the fact that patients treated with TRT often present
with disseminated disease, and after prior treatment with surgery, external
beam radiotherapy and/or chemotherapy, are likely to respond differently to similar
treatment with radionuclides. Consideration of radiobiological principles is
essential to enable treatment optimisation in external beam radiotherapy.
Radiobiological principles that apply to external beam
radiotherapy may be applied to TRT with suitable modifications. Thus, the
standard model of cell survival gives the fraction of cells surviving the
irradiation (SF) as a function of the dose delivered (D):
where α and β are disease- or even
patient-specific parameters related to radiosensitivity, and the ratio of these
parameters determine the shape of the cell survival curve (Figure 2). It is
considered that a governs cell death
from single hits, whilst b is dependent
on the absorbed dose rate. It is therefore this term that is of greater
importance in TRT . A dose protraction factor, G, has been added to this
model [22, 23] to accommodate the effect on cell kill by the change in absorbed
dose rate, resulting in the more precise equation:
where μ is the constant of sub-lethal damage repair and
t� is a time-point during the treatment prior to time t.
In practice, the large variation in absorbed dose rates for
the radionuclides used in TRT mean that the application of this model is
somewhat impractical. The Biologically Effective Dose (BED) was introduced to
address these concerns [24, 25] and is defined as
This model has been used to compare absorbed doses delivered
with TRT, with those delivered with external beam radiotherapy using the
For external beam radiotherapy:
and for TRT:
As TRT is frequently used to treat patients with a wider
variation in disease progression and treatment background, and because of the
implications of the heterogeneity of uptake of a radionuclide , the
application of radiobiological concepts are arguably of greater relevance than
is the case for external beam radiotherapy, although this remains an area in
need of significant research . It is possible that radiobiological
arguments may be employed to combine TRT and external beam radiotherapy .
Status of dose calculational approaches
Dose calculations for diagnostic agents are developed using
a combination of animal data and human subject (healthy volunteers or patients)
during the drug approval process. Animal data may be extrapolated by a variety
of methods, none of which is necessarily more correct than the other .
Human data are most often analysed using the conjugate-view approach described
above, although for some positron emitting agents, positron emission tomography
(PET) may be used to obtain quantitative data for dosimetry . Dosimetry for
these agents developed as typical values for average adults and children [31,
32] are usually accepted as adequate
Imaging of patients to obtain anatomical and physiological
information has progressed substantially in the last decade. Anatomic
information obtained from medical images, e.g. with MRI or CT, can be expressed
in 3 dimensions (3D) in voxel format, with typical resolutions on the order of
1 mm. Similarly, SPECT and PET imaging systems can provide 3D representation of
activity distributions within patients, with typical resolutions of around 5-10
mm. The newest systems now combine CT with both PET and SPECT state of the art
imaging systems on the same imaging gantry, so that patient anatomy and tracer
distribution can be imaged during a single imaging session without the need to
move the patient, thus greatly improving and facilitating image registration.
The use of Monte Carlo radiation transport codes with knowledge of patient
anatomy will result in a significant improvement in the accuracy of dose
calculations. Radiation dose calculations for nuclear medicine applications
have been mostly relegated to abstract and theoretical calculations, used to
establish dosimetry for new agents and provide reasonable dose estimates to
support radiopharmaceutical package inserts and for use in open literature
publications. When patients are treated in therapy with radiopharmaceuticals,
careful, patient-specific optimisation is not performed, as is routine in
radiation therapy with external sources of radiation i.e., radiation producing
machines and brachytherapy. There are several reasons for this. One involves
the limitations on spatial resolution and accuracy of activity quantification
with nuclear medicine cameras. Another has to do with the realism and
specificity to an individual patient of available body models. The models
described above were designed to represent the �reference� adult male and
female, children, and so on. Besides using geometric primitives to represent
the body and its various organs, only one model is available for any category
of individual, so dose estimates calculated using this approach will contain
significant uncertainties when applied to any subject, and physicians
understandably have low confidence in the use of these results to plan
individual subject therapy. Thus, unfortunately for the patients, a �one dose
fits all� approach to therapy is usually employed, with significant caution
resulting in administration of lower than optimum levels of activity to the
majority of subjects. The use of image-based models, not only to develop new
�reference� phantoms, but also to permit the use of patient-specific models for
each therapy patient, is now well-developed. Internal dosimetry is thus poised
to truly enter a �Golden Age� in which it will become a more integral part of
cancer care, as dosimetry is used in external source radiotherapy. The realism
of the newer models is shown in Figure 3, with comparison to the form of the
existing models developed and implemented in the historical MIRD system.
Image-based Computational Tools
Several efforts to use image data in dose calculations, as
described above, include the 3D-ID code from the Memorial Sloan-Kettering
Cancer Center , the SIMDOS code from the University of Lund , the RTDS
code at the City of Hope Medical Center , the RMDP code from the Royal
Marsden Hospital  and the DOSE3D code . The code with the most clinical
experience to date is the 3D-ID code. These codes either rely on the standard
geometrical phantoms (MABDose and DOSE3D) or patient-specific voxel phantom
data (3DID and SIMDOS) and various in-house written routines to perform photon
transport. Neither has a particularly robust and well-supported electron
transport code, such as is available in EGS ,
MCNP , or GEANT . The PEREGRINE code 
has also been proposed for three-dimensional, computational dosimetry and
treatment planning in radioimmunotherapy.
The usual approach used in these codes is to assume that
electron energy is absorbed wherever the electron is first produced. The
development and support of electron transport methods is quite complex, as
evidenced by ongoing intensive efforts by both the EGS4 and MCNP computer code
working groups. It is not reasonable to expect in-house written codes to deal
effectively with electron transport. In areas of highly non-uniform activity
distribution, such as an organ with multiple tumours with evidence of enhanced
uptake of an antibody, explicit transport of both photons and electrons is
needed to characterise dose distributions adequately.
Clinical experience with dosimetry
Clinical applications of dosimetry to TRT have to date been
limited to sporadic instances rather than comprehensive clinical trials. The
lack of standardised methodology means that results are difficult to compare
directly, although there is emerging evidence that dosimetry can prove to be of
practical benefit in aiding the clinician to produce informed decisions.
Although dosimetry has been applied to a wide variety of diseases, there are a
number of specific examples where it is likely to be of particular benefit.
Benign thyroid disease
There has been continual debate on the potential application
of dosimetry to benign thyroid disease. Several authors have advocated that
administered activities should be tailored to the individual patient to deliver
a prescribed absorbed dose, although the majority of therapies are based on
fixed activities. However, it has been shown that absolute uptake of
radioiodine varies widely from patient to patient and is markedly more
pronounced for autonomous nodules than for normal tissue . The same authors
also examined the use of I-124 NaI PET to perform tracer dosimetry and
concluded that absorbed dose estimates could be made with an accuracy of within
10%. It has been shown that an absorbed therapeutic dose can be predicted by a
prior tracer administration to within a degree of accuracy that would enable
patient-specific treatment planning [45, 46]. It has been further
shown that the rate of hypothyroidism resulting from the treatment of Graves�
disease with radioiodine is correlated with the absorbed dose .
Treatment of thyroid cancer with 131I NaI is the
most common oncological application of TRT and has been used for nearly seven
decades. There have been few changes in treatment regimens in that time,
although there is no internationally agreed standard on how to perform
treatment. In the majority of cases, treatments are based on fixed activities
rather than absorbed doses, although there have been exceptions [48, 49].
Typically, patients will be given 1-3 GBq for ablation, and 3-20 GBq for a
subsequent therapy . Benua et al.  have administered according to a
whole-body dose. Some authors have used 124I NaI to perform
dosimetry for the treatment of thyroid cancer [52, 53]. Where dosimetry has
been performed, there is ample evidence that a wide range of tumour absorbed
doses are delivered from fixed activities [53, 54]. It is probable that because
of the relative simplicity of treatment, the lack of complications caused by
other therapies, and its widespread use, radioiodine treatment of thyroid
cancer offers the greatest potential for determining the dose-response criteria
in a multi-centre setting that would lead to patient-specific treatment .
The issue of thyroid stunning would need to be circumvented although there is
still doubt as to the level of diagnostic activity above which stunning occurs,
whether the stunning phenomena is relevant for 123I, 124I,
and indeed whether stunning is a real effect as such or whether it is an early
therapeutic effect [56-58] .
MIBG therapy for neuroendocrine tumours
I-131 meta-iodobenzylguanidine (mIBG) has been used for 20
years for the treatment of adult and paediatric neuroendocrine tumours, including
phaeochromocytoma, paraganglioma and neuroblastoma. Administration protocols
vary widely from standard administrations of 7.4 GBq to activities larger than
30 GBq for adults [59-61] . As with radioiodine treatment of
thyroid cancer, the number and frequency of administrations also vary widely
from centre to centre. Current problems with this therapy that could be
addressed with dosimetry include the issue of carrier-added mIBG, since at
present only a small fraction of the mIBG that is infused is labelled with
I-131. Where dosimetry has been performed it has been shown that a wide
variation in absorbed doses to either the whole-body or to the tumour or normal
organs result from fixed administered activities [60, 62, 63] . A
multi-centre dosimetry-led clinical trial aimed at relapsed or refractory
neuroblastoma that recently commenced in Europe aims to administer a whole-body
absorbed dose of 4 Gy in 2 fractions . This and similar trials in the US
are leading to the delivery of relatively high activities.
The use of monoclonal antibodies for cancer treatment has
been well established and is currently an area of clinical research that is
rapidly increasing. The most common target for monoclonal antibody (mAB)
therapy is lymphoma, with a number of centres having developed their own mAB�s
[64-66] . Two products, Bexxar and
Zevalin have been approved by the US FDA for treatment of relapsed or
refractory B-cell non-Hodgkin�s lymphoma. Both use the anti-CD20 antibody
although the major difference is that Bexxar employs 131I as the
radionuclide whereas Zevalin uses the longer range beta emitter 90Y.
The clinical efficacies of these treatments have yet to be directly compared in
a trial. A key element of the treatment with Bexxar is that administration is
based on a whole-body dose of 0.75 Gy , whereas for Zevalin, dosimetry was
not recommended .
Peptide therapy for neuroendocrine tumours has recently
emerged with the development of somatostatin analogues such as the compound DOTA-DPhe(1)-Tyr(3)-octreotide
(DOTATOC). Although relatively few centers have performed this treatment, there
have been a small number of dosimetric studies carried out . Barone et al.
 measured the kidney absorbed dose from administrations of 8.1 GBq-22.9 GBq
of 90Y DOTATOC, and after taking into account the biologically
effective dose (BED), found a strong correlation between BED and creatinine
clearance. A study by Hindorf et al (in press) found that tumour absorbed doses
resulting from a fixed administration of 90Y DOTATOC varied widely
on an inter-patient basis, although in repeated treatments the intra-patient
variation was much smaller, indicating that it would be possible in principle
to use dosimetric results from the first therapy to adjust subsequent
therapies. Studies are ongoing to compare the relative efficacies of DOTATOC
with DOTATATE and the optimal radionuclide .
Summary of Clinical Experience
The clinical introduction of internal dosimetry for TRT has
been slow and is still far from being implemented routinely, despite a European
directive stating that 'For all medical exposure of individuals for
radiotherapeutic purposes, including nuclear medicine for therapeutic purposes,
exposures of target volumes shall be individually planned; taking into account
that doses of non-target volumes and tissues shall be as low as reasonably
achievable and consistent with the intended radiotherapeutic purpose of the
exposure' . In cases where dosimetry is performed, the methodology
employed is adapted from calculations derived for radiation protection rather
than radiotherapy. Recent studies have shown conclusively that the
administration of fixed activities results in a wide range of absorbed doses
and there is now initial evidence to suggest that patient outcome is more
likely to be correlated with absorbed dose rather than administered activity.
It is likely that in the near future, internal dosimetry for both tumour and
normal organs will become routine clinical practice, aided by improved
techniques and possibly the application of radiobiological considerations. This
will facilitate individualised treatment planning and the administration of
cocktails of radionuclides.
The case for optimisation
Physicians generally administer similar levels of activity
or activity per unit total body mass to all patients . This has been
reasonably successful in the use of radioiodines against thyroid cancer and
hyperthyroidism, as the �therapeutic window� (difference in dose levels between
what is experienced by the tumour and that experienced by the most important
normal tissue) is large. Nonetheless, some centres are now moving towards the
use of patient-specific dose calculations even for iodine therapy, to optimise
the therapeutic regime and to try to minimise the risk of unwanted side effects
such as sialadenitis and sicca syndrome . In other recently evolving forms
of therapy, however, (e.g. the use of monoclonal antibodies and radiolabeled
peptides in therapy), the tumour-to-normal tissue absorbed dose ratio may be
low. Without the use of a patient-specific treatment planning strategy based on
radiation absorbed dose, patients are frequently treated cautiously and given
low amounts of the therapeutic agent, to avoid deleterious effects in normal
tissues (most notably the bone marrow). Different patients will have different
levels of tumour and normal tissue uptake concentrations, as well as in the
clearance rates at which activity leaves these tissues. Patients who clear the
activity more slowly from their bodies will necessarily receive higher doses to
marrow and other normal tissues than those with faster rates of elimination.
Thus, only some patients will receive optimal care, and a majority of patients
will receive a lower than optimal administration of activity. This usually
results in no deleterious effects in normal tissues, but suboptimal therapy
being delivered to the malignant tissues, with poor response rates and high
rates of relapse. As was stated by Siegel et al. :
�If one were to approach the radiation oncologist or medical
physicist in an external beam therapy program and suggest that all patients
with a certain type of cancer should receive the exact same protocol (beam
type, energy, beam exposure time, geometry, etc.), the idea would certainly be
rejected as not being in the best interests of the patient. Instead, a
patient-specific treatment plan would be implemented in which treatment times
are varied to deliver the same radiation dose to all patients. Patient-specific
calculations of doses delivered to tumours and normal tissues have been routine
in external beam radiotherapy and brachytherapy for decades. The routine use of
a fixed GBq/kg, GBq/m2, or simply GBq, administration of
radionuclides for therapy is equivalent to treating all patients in external
beam radiotherapy with the same protocol. Varying the treatment time to result
in equal absorbed dose for external beam radiotherapy is equivalent to
accounting for the known variation in patients� uptake and retention half-time
of activity of radionuclides to achieve equal tumour absorbed dose for
internal-emitter radiotherapy. It has been suggested that fixed
activity-administration protocol designs provide little useful information
about the variability among patients relative to the normal organ dose than can
be tolerated without dose-limiting toxicity compared to radiation dose-driven
Thierens et al.  noted that ��patient-specific dose
calculations in radionuclide therapy are difficult to perform and possibly
subject to large error. Therefore, individual dosimetry-based activity
calculations are not routinely applied yet and a large variety of methodologies
exists for determining the administered activity in clinical practice�� They
also noted, however, that ��as absorbed dose estimates become more
patient-specific, an improved correlation between the administered activity and
the clinical outcome may be expected. It is clear that a patient-specific
treatment planning will improve the quality of radionuclide therapy
substantially, especially in a curative setting.�
Treating all nuclear medicine patients with a single,
uniform method of activity administration amounts to consciously choosing a
lower standard of care than patients who receive radiation externally for
cancer treatments. Some have insisted that hypothesis-driven testing proves
statistically that nuclear medicine therapy patients treated with consideration
of individual dosimetry have better and more durable outcomes than those treated
under the current practice of administering the same activity levels to all
patients, as a prerequisite to considering dosimetry as routine practice. This
sets a very high hurdle for the inclusion of this practice, and certainly puts
it off many years into the future. In the meantime, tens of thousands of
patients will be receiving suboptimal therapy, and no widespread data gathering
will occur to improve dosimetry methods and understanding of the relationship
between doses received and outcomes observed. We take it as given that response
to radiation will correlate with absorbed dose more closely than it will with
administered activity, or the intention to treat. It is essential that in all
forms of radiotherapy with internal emitters, a patient-individualised dose
calculation be made when possible for the most important tumours for which a
specific uptake of the radiopharmaceutical can be derived, and for the most
important normal tissue at risk (generally the bone marrow, but possibly the
lungs, kidneys, or other organs). This is needed not only to provide a better
quality of therapy to patients treated currently, but also to establish a
database of literature that can be used to understand the variability between
subjects and the range of uptake and clearance values to be expected for
different therapy agents. Standardised methods for calculating dose are well
established and automated at present, and should be used to provide dose
calculations that are comparable and reproducible between institutions.
Figure 1 General time/activity curve for an internal emitter.
Figure 2 Cell survival curves as a function of the a/� ratio.
Figure 3 Comparison of the realism of the traditional MIRD body models with those being used to support current dose modelling efforts.
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|Received 12 January 2007; accepted 24 January 2007
Correspondence: Department of Radiology and Radiological Sciences, Vanderbilt University, 1161 21st Avenue South, Nashville, TN 37232-2675, United States. Tel.: (615) 343 0068; Fax: (615) 322 3764; E-mail: firstname.lastname@example.org (Michael G. Stabin).
Please cite as: Stabin MG, Flux GD,
Internal dosimetry as a tool for radiation protection of the patient in nuclear medicine, Biomed Imaging Interv J 2007; 3(2):e28
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