Download positron emission tomography

POSITRON EMISSION TOMOGRAPHY
A.M.J. Paans
PET-Center Groningen University Hospital, Groningen, the Netherlands
Abstract
Positron Emission Tomography (PET) is a method for determining
biochemical and physiological processes in vivo in a quantitative way by
using radiopharmaceuticals labeled with positron emitting radionuclides
as 11C, 13N, 15O and 18F and by measuring the annihilation radiation
using a coincidence technique. This includes also the measurement of the
pharmacokinetics of labeled drugs and the measurement of the effects of
drugs on metabolism. Also deviations of normal metabolism can be
measured and insight in biological processes responsible for diseases can
be obtained.
1.
General introduction
The idea of in vivo measurement of biological and/or biochemical processes was already
envisaged in the 1930's when the first artificially produced radionuclides, which decay under
emission of externally detectable radiation, of the biological important elements carbon, nitrogen
and oxygen were discovered with help of the then recently developed cyclotron. These
radionuclides decay by pure positron emission and the annihilation of positron and electron
results in two 511 keV γ-quanta under a relative angle of 180o which are then measured in
coincidence. This idea of PET could only be realized when the inorganic scintillation detectors
for the detection of γ-radiation, the electronics for coincidence measurements and the computer
capacity for data acquisition and image reconstruction became available. For this reason Positron
Emission Tomography is a rather recent development in functional in vivo imaging.
PET employs mainly short-lived positron emitting radiopharmaceuticals. The radionuclides
employed most widely are: 11C (t½ = 20 min), 13N (t½ = 10 min), 15O (t½ = 2 min) and 18F (t½ =
110 min). Carbon, oxygen, nitrogen and hydrogen are the elements of life and the building stones
of nearly every molecule of biological importance. However, hydrogen has no radioactive isotope
decaying with emission of radiation which can be detected outside the human body. For this
reason a fluorine isotope is often used as a replacement for a hydrogen atom in a molecule. Due
to these short half-lives the radionuclides have to be produced in house, preferably with a small,
dedicated cyclotron. Since the chemical form of the produced radionuclides can only be simple,
input from organic- and radiochemistry is essential for synthesis of the desired complex
molecule. Input from pharmacy is required for the final formulation and pharmacokinetic studies
and medical input is evident and required for application. Longer lived positron emitting
radionuclides are sometimes commercially available or obtainable from research facilities with
larger accelerators. Some examples of longer liver positron emitting radionuclides are 52Fe (t½ =
8.3 h), 55Co (t½ =17.5 h) and 124I (t½ = 4.2 d). Sometimes also positron emitting radionuclides
can be obtained from a generator system. Examples are 82Rb (t½ = 76 s) from 82Sr (t½ = 25.5 d)
and 68Ga (t½ = 68 m) from 68Ge (t½ = 270 d). Although all these radionuclides are used, the
isotopes of the biological most important elements receive most attention.
At the moment small dedicated cyclotrons are a commercially available product. These
1
accelerators are one or two particle machines with fixed energies. At the moment mostly
negative-ion machine are being installed because of their relative simple extraction system and
high extraction efficiency. They are installed complete with the targetry for making the four
above mentioned short-lived radionuclides in batches up to 100 GBq or more. Also the chemistry
for some simple chemical products is incorporated e.g. 11CO2, 11CO, C15O, C15O2, H215O etc..
Sometimes more complex syntheses, e.g. 18FDG, 18F-DOPA, H11CN, 11CH4 or 13NH3, are also
available from the cyclotron manufacturer or a separate specialized company. These products
become available via dedicated, automated systems or via a programmable robotic system. Other
radiopharmaceuticals have to be set up individually in each PET center.
The state of the art positron camera is a complex radiation detection technology product
combined with a relative large computing power for data acquisition and image reconstruction.
The basic detector in a modern PET camera is a BGO detector block divided in e.g. 8x8 subdetectors and read out by 4 photomultiplier tubes (PMT). By adding and subtracting the
individual signals of the PMT's the scintillating sub-detector in the BGO block can be identified.
Around 70 blocks will form a ring and 4 of these rings can added to get an axial field of view of
approximately 15-16 cm. In this way 31-63 planes are imaged simultaneously with a spatial
resolution of 4-7 mm FWHM depending on the specific design of the tomograph. The septa
between the adjacent sub-detector rings can also be retracted creating a much higher sensitivity in
this 3-D mode at the cost of a larger scatter fraction. With the present generation of positron
camera=s the singles count rates that can be managed are in the order of over 50,000,000 counts
per second resulting in coincidence count rates of over 500,000 per second. Hardware and
software for data acquisition, image reconstruction and for image manipulation is available.
Positron cameras are able to measure the radioactivity in absolute terms, Bq/pixel, which is an
unique feature. This is possible because the coincidence technique allows for the correction of the
attenuation of radiation inside the body of the individual patient. This correction is accomplished
by making an individual "transmission image" with an external positron emitting source. This
individual transmission image can also been used to correct for scattered radiation present in the
image after a 3D-acquisition. This external source is built into the camera and can be extended
from its well shielded storage box during the operation of the positron camera. To translate the
measured radioactivity distribution into functional or physiological parameters, compartimental
models have been developed for radiopharmaceuticals with known metabolite profiles. Although
only a few measurable quantities, i.e. tissue and plasma concentration (the latter by taking blood
samples), are available, it is still possible to calculate e.g. the glucose consumption by employing
a dynamic data acquisition protocol in combination with a compartimental model. It is also
possible to make a whole body scan by translating the patient through the PET camera. By
projection the transverse section images a whole body overview can be made.
A PET center is the combined relevant knowledge of chemistry, medicine, pharmacy and physics
and a PET center is staffed by all these disciplines in a good cooperating team.
2.
Accelerators for PET radionuclide production
In the energy range from 10-20 MeV all four basic radionuclides, 11C, 13N, 15O and 18F, can be
produced. In general the choice is to use a cyclotron, not a linear accelerator, because in this
energy range the cyclotron is a versatile and economic solution. Although the higher the energy
the more of the excitation can be exploited and the higher the yield, some companies on purpose
designed cyclotrons at the low energy range of 10-11 MeV protons for economical reasons.
2
Possible commercial PET cyclotron manufacturers are CTI (USA), Ebco Industries (Canada),
General Electric (USA), IBA (Belgium) or Oxford Industries (UK). In the following a more
general approach for radionuclide production is taken but at the end the focus is again on the
details of the production four basic PET radionuclides.
2.1
General production formulae
The production of radionuclides can be achieved by using neutrons or by using charged particles
as irradiation source. Irradiation by neutrons leads to neutron capture and so to neutron rich
nuclides. The use of charged particles like protons, deuterons, helium-3 or helium-4 leads to
nuclear reactions of the type (p,xn) or (p,α) which results in the production of neutron deficient
nuclides. Both types of nuclear reactions lead to radioactive radionuclei and the farther away
from the line of stability the shorter the half life will be in general.
The energy requirement of a good yield for charged particle induced reactions will be roughly
equal to the mass difference with 5 - 10 MeV added to be well above the threshold energy and to
reach the maximum in the reaction cross section.
Nuclear reactions induced by electrons have a very low cross section because of the weak
interaction of electrons with matter. In case of electrons mostly the "bremsstrahlung" generated
on a heavy target is used. So we have γ-ray induced reactions which have a very poor selectivity
for a nuclear reaction channel. Due to this poor selectivity the isolation of the desired
radionuclide often requires quite some chemistry to be performed.
Production by charged particles is controlled by following:
dNf = Ni Nt σ dt - λ Nf dt
with
Nf = number of nuclides produced
Ni = number of incoming particles
σ = partial reaction cross section
λ = ln2/t½ = decay constant
Nt = number or target nuclei
The first term gives the production rate while the latter gives the loss by decay during the
irradiation. Integration leads to:
Nf (t) = Ni Nt σ (1 - e-λt)/λ
With
Ni = i/(Zi e)
i = beam current
Zi = charge of incoming particle
e = elementary charge
Nt = m NA / M
m = weight in g/cm2
M = molecular weight
NA = Avogadro's number
3
Since σ is a function of energy
dNf (t,E) = (NA i)/(Zi e M) dm/dE (1 - e-λt)/λ σ(E) dE
with dE/dm = dE/d(ρx) = stopping power (Bethe)
Integration over the energy range of interest leads to:
Nf (t,E) = (NA i)/(Zi e M) (1 - e-λt)/λ f (dE/d(ρx))-1 σ(E) dE
So the yield is current and not time determined. Irradiation times of more than two half lives are
not productive. For radionuclide production accelerators with a high beam current and larger
beam size on the target position to avoid heating problems are required. This is just the opposite
of what experimental nuclear physics generally requires.
After irradiation the normal laws of radioactive decay are valid:
Nf (t) = Nf(EOB) e-λ(t-EOB)
with EOB = End Of Bombardment
The relation between the activity in Bq and the total number of radionuclei involved can be
calculated according to:
A(t) = N(t) - N(t+1) = N(t) (1 - e-λ) if λ<<1 s-1 the relation becomes:
A(t) = λ N(t)
In case of a radioactive chain one has to do also with "daughter" nuclides. The number of
daughter nuclides is given by the so called "Mother-daughter" relationship:
N2(t) = N2(0) e-λ2 t + N1(0) {λ1/(λ2 - λ1)} {e-λ1 t - e-λ2 t}
For the ideal nuclear reaction leading to a high yield a large cross section for the nuclear reaction
involved and a low stopping power for the particles and target material used, is required. In the
next table an example of stopping powers for different particles at different energies for
aluminum are given.
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Table 1. Stopping power (dE/d(ρx) = MeV cm2/g) for aluminium for different particles at
different energies
p
d
3
He
4
He
20
40
60
80 MeV
19.62
33.83
184.9
229.1
11.38
19.68
108.3
135.2
8.34
14.34
79.02
98.87
6.72
11.48
63.17
79.09
So the stopping power is increasing with increasing atomic number of the incoming particles and
is decreasing with increasing energy. Using heavy ions (A>4) for radionuclide production will be
hampered two fold: i) low cross sections and ii) relative high stopping powers. Conclusion: if
possible use protons or deuterons.
Fig 1. Example of excitation functions: 76Se(p,xn)77-xBr nuclear reactions. Solid lines:
experimental excitation curves for (p,xn) on 76Se. Dashes lines: theoretical excitation function
according to the ALICE code.
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The excitation functions of (p/d/τ/α,xn) reactions gives the cross section as function of the
incoming particle energy. There will always be an overlap in reaction channels: (p,n+1) starts
while (p,n) still continues often as an evaporation reaction, at higher energy direct mechanism
play a more important role. To make an evaluation of the cross sections involved for nuclear
reactions nuclear evaporation codes can be used, e.g. the code ALICE available from the Nuclear
Energy Agency in France.
2.2
Accelerators and specific activity
With charged particles neutron deficient radionuclides are produced. In a nuclear reactor neutron
capture is the most important nuclear reaction leading to neutron rich radionuclides. So both
production possibilities are complementary. There are very few overlaps. An interesting overlap
example is the production of 18F by starting with neutrons to generate tritons which induce a
charged particle reaction:
In a nuclear reactor: 6LiCO3 + n Æ t + 16O Æ
18
F+n
By the irradiation with neutrons tritons are generated by the breaking up of the 6Li into a tritium
and a helium-3 nucleus. The triton is than able to produce 18F from the 16O nucleus. Because of
the incorporation of the oxygen inside the molecule, the range of the tritons is not essential. Since
this is a two-stage process the yield is lower than the yields achievable with direct reactions as
can be done with charge particle induced reaction like:
p + 18O Æ
18
F+n
The beam energy required for (p,xn) reaction is a rule of the thumb: 7 + 10*xn MeV
with xn the number of neutrons to be knocked out or evaporated. For an exact calculation one
should calculate the mass difference and add roughly 7 MeV for the maximum cross section.
For radionuclide production the beam quality (dE/E) is not that important. The beam size should
be not to small in order not to have to high power density on the target which could initiate
problems like melting or evaporation, so a beam size of cm2 instead of mm2 has to be preferred.
The targets can be installed inside or outside the cyclotron. With internal targets the extraction of
the beam is avoided, so beam losses are also avoided. A higher yield with respect to external
targets has to be expected. At negative ion cyclotrons the extraction efficiency can be very close
to 100% due to the effect that by stripping the electrons a natural extraction becomes true. In case
of target problems it is more easy to process the target at an external target position than it is with
the target at an internal position. Also a radioactive contamination of the cyclotron, due to an
internal target, can interfere with the normal maintenance program.
The transport of the irradiated material can be very easy in the case of a gas target. Just a normal
flow can cary the radioactivity over rather long distances. In case of fluid target a helium flow
through thin tubing can push the irradiated material into the wished position. With solid target an
exchange system or train system can transport the target or target material. The local situation
will dictate the solution of the particular problems.
Classes of accelerators for radionuclide production
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Ep < 20 MeV (p,n), (p,α), (d,n), particles: p, d. Typical cyclotron for PET-Centers
Ep < 35 MeV up to (p,3n), as a multi-particle (p, d, τ, α) cyclotron this is a versatile radionuclide
production machine. Users are the commercial radionuclide producers.
Ep < 70 MeV up to (p,5n). Often a multi-particle, variable energy cyclotron. A complex
production machine. Often build as a research machine for nuclear
physics where radionuclide production is only minor interest. Examples
are the former cyclotron at the KVI (Groningen, Netherlands) and
cyclotrons at PSI (Zürich, Switzerland). Radionuclides like 123I,
81
Rb,82Sr/82Rb, 52Fe, combined productions 52Fe + 55Co can be performed
with these cyclotrons.
Ep > 100 MeV Examples are linear accelerators at BNL (Brookhaven, USA) and LANL (Los
Alamos, USA). These are research accelerators used for radionuclide
production for scientific goals primarily. Due to the high energy,
spallation reactions are the main reaction mechanism.
The technical staff required for these different classes of accelerators show of course a
tremendous differences.
Recent developments in cyclotrons for radionuclide production are the negative ion machines: H-,
D- is accelerated. To have a high beam current an external source with axial injection is required.
The big advantage of negative ions is that after passing the through a carbon stripper foil all
electrons are removed and beam is automatically bended outwards the machine. An extraction
efficiency of 100% is possible and by positioning the stripper foil only half way the beam also
multiple (at least 2) extracted beams are possible. Commercial negative ion machine are available
in wide energy range, 10-230 MeV. The demands on the vacuum during acceleration are higher
than for postive ion machines because by stripping off one electron by the residual gas not only
the acceleration process is stopped but the resulting energetic neutral beam can damage the
cyclotron. If more than one electron is stripped the beam will bend into the opposite direction and
hit the vacuum chamber or internal parts of the cyclotron.
At the moment there are also developments in low energy (8 - 12 MeV) RFQ-Linacs for the
accelaration of 3He beams in combination with special target to produce the required PET
radionuclides.
Due to the nature of the induced nuclear reaction (p,xn) here is a change in element:
Z Æ Z+1
The term "carrier free" is used when no cold material of the same chemical as the radioactive is
present. This is very difficult because often a natural dilution is the case. Often the term "non
carrier added" or "nca" production is used. This means no cold material of the same chemical
identity is added on purpose during the preparation of the radiopharmaceutical.
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Specific activity is the amount of activity per gram or mole. The theoretical maximum of the
specific activity a few radionuclides:
11
C
14
C
123
I
9.2 * 109 Ci/mol = 340 TBq/mmol
6.2 * 101 Ci/mol = 2.3 MBq/mmol
2.4 * 108 Ci/mol = 8.9 TBq/mmol
(5730 yr vs 20.4 min)
The maximum theoretical specific activity is determined by the half life. In practise these
theoretical maxima in specific activity are never reached. Very special precautions have to be
taken to keep the dilution factor within bounds. With a carrier free or non carrier added synthesis
one expects no toxic effects (e.g. H11CN is not toxic any more) and no physiological effects
(tracer principle).
2.3
Targetry and specific PET productions
The produced radionuclide can only be obtained in a simple chemical form. Irradiating complex
chemical structures causes problems because of the large energy deposition inside the target
material which can damage the chemical structure of molecule. In nuclear reactions energies of
MeV's are required. For the chemical binding in molecules energies in the order of eV are
required.
2.3.1 Targetry general
Gas targets, liquid targets or solid state targets can be used. A general problem is the cooling of
dissipated power ∆E.i. The 11C-production with a 15 MeV proton beam at 30 µA generates 450
W. One of the most critical parameters for the life time of a target is the power density (W/cm2)
which in fact is gouverned by the beam size. Increase the beam size in order to decrease the
power denstity is often advisable. Accidents, because of malfunctioning of the target, happen, in
most cases, rather soon after the start of the irradiation. Water cooling on the back-side and
helium-cooling on the entrance foil, double foil technique, in a closed circuit should be used. The
energy loss in entrance foil should as small as possible. The mechanical properties of the foil
should be able to hold the pressure from inside the target and should have a good heat
conductivity in order to get rid of the energy deposition of the beam. The energy loss in target
material should be optimized based on excitation function The low energy part of the beam, after
transmission through the active part of the target, can be dumped in the back side of target. In this
way activation of the cooling water is inhibited.
Sometimes one can combine two targets behind each other and so use the whole energy range of
the beam. An example is the simultaneous production of 52Fe and 55Co by solid target
combination: 55Mn(p,4n)52Fe and 56Fe(p,2n)55Co. These are also examples of target materials
with a high melting point. Melting and/or evaporation of target material can be serious problem,
e.g. pure selenium as target material for the production of bromine will cause problems because
of the evaporation of the selenium. Selenium copper or silver alloys have high melting points (
1000 0C or more) instead of an evaporation temperature of just over 200 0C as is the case for pure
selenium.
All materials which can be hit by the beam should be selected for a minimal production of longer
lived radionuclides. So aluminum or copper as target holder and e.g. graphite to stop the beam.
8
Avoid iron or stainless steel because many long living radionuclides can be produced. However,
since the entrance foils has to be thin and strong, e.g. havar, stainless steel or titanium foils have
to be used, thickness 8-25 µm, because of their strength. A helium cooling of the foils, by using a
double entrance foil technique can be very effective. A better cooling of the target is possible by
rotating the target. Since in fact a duty cycle of less than 100% is introduced herewith, it is not
effective for the production. By positioning the target in an inclined position, instead of
perpendicular to the beam, also a lower power density can be arranged.
There are always different nuclear reactions possible to produce a radionuclide, an example for
11
C is given in table 2.
Table 2. Nuclear reactions for the production of carbon-11
Particle
Reaction
γ
p
12
d
3
He
4
He
C(γ,n)11C
11
B(p,n)11C
12
C(p,pn)11C
14
N(p,α)11C
10
B(d,n)11C
12
C(d,p2n)11C
9
Be(3He,n)11C
10
B(3He,pn)11C
11
B(3He,p2n)
12
C(3He,4He)11C
16
O(3He,24He)11C
9
Be(4He,2n)11C
10
B(4He,p2n)11C
11
B(4He,p3n)11C
12
C(4He,4Hen)11C
E(thresh)
(MeV)
18.7
3.0
20.3
3.1
0.
24.4
0.
0.
2.3
0.
6.3
18.8
27.4
42.4
24.9
σ
(mb)
4
250
100
250
250
60
50
280
30
300
50
17
50
2.3.2 11C production
The most commonly used reaction is: 14N(p,α)11C. The target material is nitrogen with a little
oxygen mixed in: N2 (99.9999%) + O2 (2%), pressure 7 - 10 bar depending on target and beam
energy.
The primary products in the target are 11CN radicals, recoil reaction of 11C with N2, and 11CO as
recoil reaction of 11C with O2. The 11CN radicals and the 11CO are than, during irradiation,
oxydized to 11CO2. The 11CO2 is collected in a liquid nitrogen trap and than over distilled into the
chemistry set-up for further chemical synthesis.
Be aware of possible chemical reactions inside the target during irradiation. For instance when
CH4 (methane) is irradiated with protons polymerization of the target gas will result in a yellow
coating of the target wall and all radioactivity is adsorbed in this material. The molecular
structure of the target material should inert with respect to possible processes induced by the
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beam.
Fig. 2 11C - target system for the proton induced reaction on nitrogen.
Fig. 3 11CO2 collection system. Gas samples can be taken for analysis. The yield can be measured
by trapping the 11CO2 in the NaOH. Using the trap in liquid nitrogen the whole yield can be
recovered and than vacuum distilled into the chemical set-up.
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2.3.3 13N production
The most commonly used reaction for the 13N - production is 14N(p,α)13N with H2O as target
material. After irradiation the 13N is available as nitrate or nitrite in the water. Distillation under
steam with Devarda's alloy yields 13NH3. Nowadays, by addition of ethanol into the target water,
there is an in-target production of 13N-ammonia. The ammonia is most often used for cardiac
flow studies.
2.3.4 15O production
The most commonly used reaction for the production of oxygen-15 is the 14N(d,n)15O reaction.
With a positive Q-Value of 3.1 MeV a 3 MeV deuteron only cyclotron is sufficient for the
production. In fact the beam energy should not become above 6.5 MeV in order to avoid
radioactive impurities. The target material is high purity nitrogen with oxygen mixed in: N2
(99.9999%) with an addition of O2 (4%) and yields 15O2. The most commonly use of oxygen-15
is for rCBF studies. This possible is two ways: i) convert according to 15O2 + C (400oC) --->
C15O2. Upon inhalation the C15O2 is converted into H215O instantaneously in the lung
enzymatically. It is also possible to convert the oxygen into water according to: 15O2 + H2 in an
oven with Pt catalysator yields a continuous stream of H215O which can be administered
intravenously.
2.3.5 18F production
The most commonly production of 18F is the 18O(p,n)18F reaction with as target material H218O,
Oxygen-18 enriched water (>90%, costs in April 2001: US$ 160.00 /ml). The 18F is available as
ion in the water. After separation of the fluorine and the water the fluorine is available for
chemistry. The water can be used again after distillation to eliminate impurities. The enrichment
grade will diminish of course by the distillation procedure. The specific activity of the fluorinated
end product can easily be better than of carbon-11 product because fluor is less abundant than e.g.
CO2.
A second method for the production of 18F is by the 20Ne(d,α)18F reaction. To the Ne-gas F2 is
added to passivate the target chamber wall. If this passivation is not done all the produced 18F is
adsorbed to the wall and can not be extracted for further chemistry. The 18F becomes available as
F2 and the relative low specific activity depends on the amount of fluorine added before
irradiation. The two different chemical forms of 18F allow for different chemical labeling
strategies.
2.4
Commonly used radiopharmaceuticals
The most commonly used radiopharmaceutical in PET-centers is 2-[18F]Fluoro-2-deoxy-Dglucose (FDG). FDG is mostly used for studies in oncology. Due to its half life (110 min) it can
be transported over a 2 h transport distance and a lot PET camera's without an in house cyclotron
are operated on FDG only in this way. Oxygen-15 labeled water, half life 2 min, is also
frequently used in research centers for brain activation studies. The use of all other
radiopharmaceutical is often locally determined by clinical or research interest.
11
18
FDG
H215O
C 15O2
CO
11
C-tyrosine
13
NH3
11
C-raclopride
18
F-DOPA
11
C-acetate
3.
Glucose analogue for studies in brain, heart and oncology. The most used
radiopharmaceutical.
Functional brain studies (rCBF)
Functional brain studies (rCBF). Is converted into water in the lung
instantly
Cerebral blood volume studies
Amino acid for brain studies and oncology
Ammonia for blood flow studies in the heart
Dopamine receptor system, Parkinson's Disease
Dopamine receptor system, Parkinson's Disease
Cardiological studies
PET scanner
3.1
Decay of neutron deficient radionuclides
There are two decay possibilities for neutron deficient radionuclides: positron emission or
electron capture (EC):
Positron decay
p Æ n + ß+ + ν
Electron capture
p + e- Æ n + ν
The energy condition for decay by positron emission is:
Q(ß+) = M(A,Z+1)c2 - M(A,Z)c2 - 2m0c2 + I
So positron decay is only possible if an energy of 2m0c2 (= 1022 keV) or more is available,
otherwise Electron Capture will happen. In practise a rather surplus in energy is required before a
large percentage of the decay goes by the positron decay channel instead of the EC channel.
With the positron decay there are two conservation laws to be obeyed: i) Conservation of energy
tells that 1022 keV is available and ii) conservation of momentum tells that at the moment of
annihilation no momentum is available so p = 0 kgm/s before and after annihilation.
The positron is slowed down in tissue, at the end of track a positronium, a hydrogen like atom, is
formed by positron and electron. Positron and electron are anti-particles and so they will
annihilate. In singlet state of the positronium a 2 quanta annihilation (mean life time 8 ns) will
occur. In the triplet state a 3 quanta annihilation (mean life time 7 µs) will take place. The triplet
state will be formed in only 0.3%. The 2-quanta annihilation shows a finite width of 0.5o around
180o in the angular correlation measurements signaling that the momentum at the moment of
annihilation is not always exact zero.
3.2 Imaging
3.2.1. In conventional nuclear medicine
In conventional nuclear medicine a gamma camera consisting of a NaI crystal, thick 3/8" or 1/2",
with photomultiplier tubes (PMT's) and a collimator in front is used for image formation. The
position of a scintillation is calculated from:
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X = Σ PX(i) L(i) / Σ L(i)
Y = Σ LY(i) L(i) / Σ L(i)
with PX(i) and PY(i) the position coordinates of PMT(i) and L(i) the amount of light received by
PMT(i) and:
Σ L(i) = Total light output = Energy of the detected quant
A collimator is a compromise between spatial resolution and efficiency. Works optimal for
gamma radiation between 100 and 200 keV. Two modes of images can be obtained: planar
images and transverse section images by rotating the camera around the patient (SPECT= Single
Photon Emission Computed Tomography). Both modes do not yield quantitative information
because a correction for the attenuation not is possible. The thickness required for the collimator
and the thickness of septa (the lead between the holes) are a function of the energy of the gamma
rays. At a gamma energy of 511 keV the wall thickness and the thickness of the collimator have
to be increased in such a way that the hole size and the spatial resolution become competitive and
the weight of the collimator is roughly 200 kg. The efficiency of the collimator has decreased
with a factor two, with respect to a general purpose collimator at 140 keV, to approximately 5x
10-5.
3.2.2. Annihilation radiation
To image the annihilation radiation profit should be taken from its unique properties: 511 keV colinear (180o) and simultaneously! So coincidence measurements and ideally TOF (Time Of
Flight) measurements should be done. For two detectors, A and B, at a distance 2d and a point
source P at distance x from the center line the difference in distance is:
PA - PB = (d+x) - (d-x) = 2x
time involved: ∆t = 2x/c with 2x = 1 mm, ∆t = 3.3 ps
Scintillation detectors with these timing properties and high sensitivity for 511 keV radiation do
not exist at the moment. Until ~1983 mostly NaI detectors and some BaF2 detectors systems have
been used. After ~1983 BGO (Bismuthgermanate Bi4Ge3O12) material in a block-detector
structure is mainly used. Originally the detector with its single PMT was determining the spatial
resolution. To improve in efficiency BGO with its density and high Z-value is being used. To
improve also in spatial resolution a gamma camera read-out with four PMT's was designed on a
single BGO crystal. Due to the thickness of the BGO detector the scintillation light is spread out
quite a bit. By cutting the BGO detector into 8x8 sub-detectors a light guiding was build in which
allowed, together with the four PMT's, for a spatial resolution basically the same as the size of the
sub-detectors. The drawback of BGO is its relative low light output (15% of the output of NaI).
In the past also Gadolinium-orthosilicate (GSO) has been applied in combination with BGO as a
dual detector on one PMT. An advantage of GSO is the higher light output, see table 3. Due to
the differences in decay time one can determine, using pulse shape discrimination, which of the
13
two detectors is responding. Recently GSO is again being used but now as an area detector with a
gamma camera logic read-out. Recently LSO or Lutetium-ortho-silicate, has been tested for
application in PET scanners. The relative high light output (75% of the output of NaI) is an
advantage. The disadvantage of a natural radioactive component in natural Lu has no
consequences as long as coincidence measurements are performed. The first PET scanners with
LSO detectors, both for small animal studies and for whole body human studies, are
commercially available now.
Table 3. Detector materials used in PET scanners
Density (g/cc)
Eff. Atomnumber
Mean Free Path (cm)
Hygroscopic
Decay time (ns)
Relative light yield
Energy resolution*
*
NaI
BGO
GSO
LSO
3.67
51
2.88
yes
230
100%
7.8%
7.13
75
1.05
no
300
15%
10.1%
6.7
59
1.43
no
56/600
25%
9.5%
7.4
66
1.1
no
40
75%
10%
NB: This energy resolution is valid for a single crystal. Is not necessarely true for a blockdetector
In PET scanners somehow a circular of hexagonal structure has to be realized in order to perform
coincidence measurements. This can also been realized by a rotating dual headed uncollimated
gamma camera system. These systems, limited in count rate because only 2 detectors are used,
and all types of different configurations NaI of BaF2 based have been used till roughly 1983 when
the BGO block-detector was introduced. The block detector is commonly divided into 8x8 subdetectors which are read-out by 4 PMT's. The size of the block detector depends on the resolution
to be achieved. Size of the sub-detector varies from 4*4 to 6.5*6.5 mm2 and the thickness is
varying between 20 and 30 mm. See fig. 4-6. In order to increase the axial length up to 4 adjacent
detector rings are assembled into a PET scanner.
14
Fig. 4. BGO Block detector with PMT's (Courtesy Siemens/CTI)
Fig. 5. Light guiding effect of the cutting in the BGO block detector (Courtesy Siemens/CTI)
15
Fig. 6. Spatial resolution of a BGO block detector with 4 PMT's (Courtesy Siemens/CTI)
3.3 Data acquisition and image reconstruction from projections
3.3.1 Data acquisition
The organization of the acquired data can be in two forms:
1)
Event-by-event or List Mode
The position of each individual annihilation pair and some type of timing information are
individually stored. Afterwards reconstruction into sinograms in a then defined dynamic study is
performed. The time frames of the study can be chosen or redefined with this form of storage.
2)
Sinogram mode
The data acquired by a PET scanner is projection data by nature since only a coincidence and no
TOF measurement is possible. The total number of annihilation events on a Line of Response
(LOR) are stored in one matrix element according to (r,θ) coordinates which is called a sinogram
because of its behaviour for rotating point source. A LOR is the line between two detectors
operated in coincidence.
Which storage mode is chosen is depending on the amount of LOR's vs the number of events.
Examples:
i)
Dual head coincidence system: Always list mode because it is a 3D-system with > 100
MLOR's and the number coincident events measured is considerably less.
ii)
2D ring system: Always sinogram mode
16
iii)
3D high resolution ring system: To be considered based on number of LOR's and
expected number of coincident event.
Measuring along a LOR is measuring a line integral projection, so the Radon transform is
measured. The Radon transform maps the data from (x,y) coordinate system into projection data
domain, (r,φ). All points on the line (LOR) are mapped onto a single point. A point in object
space will follow a sinogram in projection space (x',φ).
3.3.2 Reconstruction by Filtered Back Projection
The most commonly used method in image reconstruction is Filtered Back Projection.
The back projected image is given by: g(r) = f f(r) h(r,r') dr'
With f(r) the real radioactivity distribution and with h(r,r') the system response function or Point
Spread Function (PSF). The deconvolution is most easy to perform in Fourier space:
G(k) = F(k).H(k)
with G(k) = f g(r) exp(2πi k.r) dr
F(k) = G(k) H(k)-1
Due to noise in the original data, the limited band width in the Fourier transform, over-emphasing
of noise easily occurs. To prevent a window with a smooth cut-off should be applied e.g.:
Hanning window: W(k) = 0.5 + 0.5 cos(π k/kmax), 0 for k > kmax
kmax = 1/2d according to Nyquist sampling theorem.
3.3.3 Image reconstruction by Maximum Likelihood Expectation Maximization
Maximum Likelihood Expectation Maximization (ML-EM) is an iterative method that maximizes
the probability of the reconstructed image for a given set of measured projection data. Each
emitted photon from a pixel b (b=1, 2, ..B) in the object is detected by a detector unit d
(d=1,2,..D) with a probability p(b,d). The unknown emission density f(b) can be estimated using
the measured projection data n*(d) in detector d.
λ*(d)=Σb f(b) p(b,d) = expected number of counts in detector d
In case of a Poisson distribution the likelihood function L is:
L = Πd exp[-λ*(d)] {λ*(d)n*(d)/n*(d)!} = Πd Pn*(d) λ*(d)
with Pn*(d) de Poisson distribution and λ*(d) the expectation of n*(d)
In the iterative scheme the difference between step k and k+1 is minimized and can be used as
stop criterion. In reality it is easier to examine the logarithm:
log(L(k+1)/L(k))=log(L(k+1)) - log(L(k))
This formula can be calculated because the p(b,d)'s are known, n*(d) is the measured projection
data. For the initial values for the f(b)=s the distribution can be assumed to be uniform. In PET
17
this ML-EM scheme showed to be successful because positron emission follows Poisson
statistics. In practice the stop criterion has to bet set by evaluating the images at different
iterations. After a certain number of iterations artifacts can be generated and this stage should be
avoided.
3.3.4 Different types of PET scan
The scans which can be made in PET scanner are:
i)
Static scan: a set of transverse section images. Interpretation by visual inspection and/or
by left/right differences. Often sufficient for a clinical study.
ii)
Dynamic scan: a set of consecutive scans in time. The distribution as function of time can
be studied in imaged area. Information as function of time is essential input for the
derivation of functional parameters. Also arterial blood sampling and analysis is often
required for the quantification of a functional parameter.
iii)
Whole body scan: a set of consecutive scans over the body. By combining these scan in a
3-dimensional volume an overview of the radioactivity in the body is visualized. A whole
body scan is often used in oncological studies.
3.4
Resolution, parallax, scatter, accidental coincidences
3.4.1 Resolution and parallax
In 3D-mode special, the radiation can penetrate the block detector from a large opening angle.
Especial is scanners with smaller radius, brain systems or small animal scanners, the identification of the Depth Of Interaction (DOI) is essential to maintain the spatial resolution. In PET
scanners with block-detectors the block structure determines the resolution mainly. If the block
substructure become smaller this effect becomes more important. Decreasing the thickness is not
an option because sensitivity is important. With a sub-detector size of 2*2 mm2, as is the case in
some experimental scanners, a lot of photons will enter not perpendicular and may be half way
the crystal. To recalculate this parallax effect into the image reconstruction DOI information is
essential. DOI can be achieved by dual layers of detectors, e.g. LSO/GSO. Because of the
different response of the detectors the different detectors can be identified by pulse shape
discrimination.
The partial volume effect will always be present, a δ-function will be seen as a Gaussian profile.
So the object will be seen larger and less prominent since all counts are distributed over a larger
volume.
3.4.2 Scattered radiation
The gamma-spectrum from a patient is quite different from the spectrum from point source in air
because a lot scatter will occur inside the body. In fact most gamma quanta will be scattered and
will not be in the photo peak. The attenuation is described by: I = I0 exp(-µx)
There are three competing processes by which the gamma radiation is attenuated:
i)
ii)
Photo-electric absorption: µpe = k.ρ/A {Z4/(hν)3} with k a constant that depends on the
atomic shell involved, ρ the density, A the atomic weight and Z the atomic number of the
material
Compton scattering, the energy of the scattered photon is given by:
18
iii)
hν' = hν0 ( 1 + α −α cosθ), with α = hν0/m0c2
or: 1/hν' - 1/hν0 = (1-cosθ)/m0c2
Pair production: this is not an issue at 511 keV
Due to the scatter the energy of the scattered radiation is decreased depending on the scattering
angle. The scatter fraction can be measured by using a cylindrical phantom with a cold rod inside.
Due to the scattering wrong LOR=s will be established resulting in a background in the cold rod.
The number of counts in the cold rod in the image gives a measure for the scatter fraction. Also
the energy window used will influence scatter fraction. A narrow energy window will decrease
the scatter fraction of course but most PET scanners are using rather wide energy windows
because of the rather low energy resolution achieved by the block-detectors. Examples are:
i)
ii)
iii)
Dual head NaI, 3D, dE/E = 10%. Measurement of the photo-peak coincidences only
yields a scatter fraction of 12-15%.
Siemens Ecat 951 (1991), BGO, 2D, window 250-850 keV yields a scatter fraction of
approximately 15%.
Siemens Ecat HR+ (1997), BGO, 3D, window 350-650 keV, scatter fraction in the brain
of approximately 30% is measured, in the chest the fraction can go up to 50%.
The scattered counts are mis-positioned in the image because of the change in direction in fact a
wrong LOR is taken.
3.4.3 Accidental coincidences
The counts acquired in a PET scan can sub-typed in the true coincidences, the scattered
coincidences and the accidental or random coincidences:
N = Ntrue + Nscatter + Nacc
Presently techniques are available, based on the measured attenuation, to have an individual
scatter correction. Accidental coincidences are possible due to the use of finite time
window:
Nacc = ∆t . Ns1 . Ns2
with ∆t time window and Nsi the singles count rate in detector i
In a gamma camera based dual head coincident system the accidental coincident count rate can be
considerable because the gamma camera manufacturers have increased the singles count rate
capability of a camera considerably in the 1990's in order offer a combined SPECT/PET camera
with improved PET properties as compared to the dual headed systems designed earlier. Pulse
shortening techniques have been used in order to compete better with the ring systems. Singles
count rates up to 2 Mcps have been reported. With often used time window of 15 ns the
accidental coincident count rate can go up quite substantially, i.e.
Nacc = 15. 10-9. 2.106 . 2.106 = 60 kcps.
The true coincident count rate did increase but the accidental coincident count rate shows a
quadratic behaviour. The total singles rate in the two detector is 4 Mcps and Nacc can go up
19
dramatically while Ntrue is limited to approximately 5 kcps. There will be a considerable
background in these images. In areas where the target non-target ratio is not that high the clinical
values of these images may be hampered.
The maximum singles count rate in block detector is approximately 250 kcps. The time window
used in a ring system is 10 ns or somewhat less. For two blocks in BGO system the accidental
coincident count rate amounts than to:
Nacc = 10.10-9 . 250.103 . 250.103 = 625 cps
Total singles rate can go up to approximately 70 Mcps and the sum of the true and scattered
coincident count rate will be over 500,000 cps. Due to the fact that ring systems have a large
number of coincidence channels, and not only one as is the case in dual headed gamma camera
based system, the accidental count rate is rather limited. Looking at a system with 288 BGO
block-detectors is in fact looking at 288 gamma cameras operated in coincidence.
The Noise Equivalent Count rate (NEC) is defined as the coincident count rate equivalent in
terms of noise in a measurement that does not include scattered or random coincidences and is
given by:
NEC = T2 / { T + 2fobjR + S }
where T, R and S are the true, random(=accidental) and scatter, and fobj is the fraction of the field
of view subtended by the object.
In fig. 7 the count rate as function of the radioactivity concentration in a 20 cm diameter 20
cm long phantom is shown in the lower diagram. In the upper the diagram the dead time as
function of the concentration is given. The maximum activity concentration to be measured
reliable is 0.025 MBq/ml. In the diagram also the delayed coincident count rate is shown. The
accidental count rate is than measured by delaying the start signal for the coincidence circuit
over more than one time window. The counts than measured are the accidental coincidences.
This data was measured on a Siemens/CTI Ecat HR+ PET scanner.
3.5 Sensitivity, attenuation correction and quantification of PET scans
3.5.1 Sensitivity
Detector sensitivity is determined by the photo-effect efficiency:
Photo-electric absorption: µpe = k.ρ/A {Z4/(hν)3} with k a constant that depends on the
atomic shell involved, ρ the density, A the atomic weight and Z the atomic number of the
material
20
Fig. 7. Count rate vs activity concentration for a standard 20 cm phantom
21
So high density and high Z-value yield a high atomic cross section. The solid angle subtended is
the other factor determining the overall efficiency. In the so called 2D mode septa, disk shaped,
are placed between the adjacent sub-detector rings. Their goal is to reduce the scattered radiation.
In the 3D mode these septa are removed so sensitivity goes up with roughly a factor of 7 because
coincidence over a large number of rings are accepted but also the scatter fraction will increase,
also because of the limited energy resolution or large energy window used, to maximal 50%. The
sensitivity profiles for the Siemens HR+ in 2d- and 3D-mode are shown in fig. 8 and fig. 9. In fig.
10 a Siemens Ecat HR+ is shown with the front panel removed.
Fig. 8. 2D sensitivity for a Siemens Ecat HR+ scanner. The 4 detector rings are visible.
22
Fig. 9. 3D sensitivity. The coincidence acceptance is over 22 sub-detector rings.
3.5.2 Attenuation correction and quantification
By using an external positron source the individual attenuation can be measured. From a blank
scan and a transmission scan an attenuation map can be calculated. Because of the coincidence
measurement the position of the source in the body is not relevant, only the total thickness.
I = I0 exp(-µ.a) . exp(-µ.b) = I0 exp(-µ.(a+b)) = I0 exp(-µ.d)
Because of the coincidence measurement the measured attenuation is indenpendent of the
position of the source and can be performed by external sources.
To speed up procedures and to gain in throughput often compromises are made in designs. To
decrease the time required and to get better statistics also a singles transmission measurement has
been proposed. Since no coincidence measurement is used this proposal has theoretical draw
backs but working is in progress to calibrate the singles attenuation as measured with a 137Cs
source at 662 keV, half life 30.2 yr, to the standard 511 keV 68Ge/68Ga annihilation source, half
life 270 d. Due to the much longer half life of the cesium source a yearly replacement as the
germanium/gallium source does is not longer required.
With the attenuation correction and the scatter correction calculated from the individual
attenuation measurement a PET scanner is able to produce quantitative images in the sense that
the activity per pixel in absolute terms can measured (Bq/pixel). This is a unique feature of PET
due to the use of coincidence measurements.
23
Fig. 10.
A Siemens Ecat HR+ positron camera with front panel removed
3.6 Possible PET-scanners anno 2001
The possible scanners anno 2001 can be divided in the following sub groups:
i)
Dual headed uncollimated rotating gamma camera's in coincidence mode
ii)
Partial ring systems
iii)
Full ring systems
iv)
PET/CT systems
v)
Experimental and high resolution scanners
3.6.1 Dual Head Coincidence Systems (NaI detectors)
All manufacturers have such a system available since their new introduction in the 1990's. The
first version of such a system was already marketed in 1968, even without computer system but
with a back-projection reconstruction in hardware. In the 1990's improved versions as a
multifunctional competitor of the ring systems were developed. Now, in 2001, with the
developments in clinical PET, these systems are not an option any more because enough patient
volume is available to operate dedicated ring systems during full working hours. The spatial
resolution of dual headed systems will be 4 to 6 mm FWHM and the maximal singles count rate
24
has been increased, values up to 2 Mcps singles rate have been reported. The "true" coincident
count rate is more difficult to obtain while the "accidental" coincident count rate is the problem as
explained before. Rotating dual head systems were a nice solution as a go between but can not
really compete with ring systems.
3.6.2 Partial and Full Ring Systems
Table 4. PET ring systems available commercially in 2001
Manufacturer
system detector
2D/3D
ADAC
C-PET Plus
Allegro
Advance NXi
mPower
Exact
ART
Accel
Exact HR+
3D Only
3D Only
2D/3D
2D/3D
2D/3D
3D Only
2D/3D
2D/3D
GE
Positron
Siemens
Attenuation correction:
Curved NaI detectors
GSO
BGO Block detectors
BGO Block detectors
BGO Block detectors
BGO Block detectors
LSO Block detectors
BGO Block detectors
68
Ge coincidence
Cs singles
Scatter correction in 3D is essential
137
Table 5. PET System parameters
Systeem
Det
Axial
(cm)
Res
(mm)
Sens-2D
Sens 3D
(kcps/kBq/ml)
C-PET+
Allegro
Advance NXi
mPower
Exact
ART
Accel
Ecat HR+
NaI
GSO
BGO
BGO
BGO
BGO
LSO
BGO
25.6(125*)
18 (90)
15.2(35)
16.6(61)
16.2(47)
16.2(47)
16.2(47)
15.5(63)
5.0
4.8
4.8
5.8
6.0
6.2
6.0
4.6
na
na
5.4
7.5
4.9
na
4.9
5.4
Companies:
ADAC:
GE:
Positron Corp:
Siemens:
C-PET+, Allegro
Advance NXi
mPower
Exact, ART, Accel, HR+
* In brackets the number of planes imaged.
25
10.8
21.6
31.1
35.1
21.1
7.3
21.1
24.3
3.6.3 Combined PET/CT systems
Table 6. Combined PET/CT systems
ADAC Gemini
GE
Discovery
Siemens
Biograph
PET: Allegro/C-PET Plus
CT: Philips Secura
PET: Advance Nxi
CT: Lightspeed
PET: HR+
CT: Somatom Emotion or Emotion Dual
Fig. 11. PET/CT system (Courtesy Siemens/CTI)
3.6.4 Experimental and High Resolution Research Scanners
i)
High Resolution Research Tomograph: LSO/GSO (Siemens/CTI). Spatial resolution 2 2.5 mm FWHM with DOI, brain system. DOI by pulse shape discrimination on LSO and
GSO signal
ii)
MicroPET: LSO with position sensitive PMT's (Concorde Microsystems). Spatial
resolution 2 mm FWHM, for rat and monkey studies.
iii)
3D HIDAC-PET: Multi Wire Proportional Counters with gamma converters (Oxford
Positron Systems). Designed for small animal studies.
26
iv)
Experimental animal PET devices at several Universities. New read-out systems: APD's
(Avalanche Photo Diodes) instead of PMT's.
v)
Dual layer PET/SPECT system with LSO/NaI (Siemens/CTI). The NaI detection system
is in use for the SPECT studies while second layer of LSO for PET will be used.
Fig. 12.
FDG brain scans made on the differente generations of PET scanners from 1975
to 1997 (Courtesy Siemens/CTI).
3.6.5 Limitations in spatial resolution
Limitations in spatial resolution for PET scanners can be divided in a fundamental and a technical
limitation. The fundamental limitation is the finite range of the emitted positron. The maximum
energy of the positron emitted is typical for each radioisotope. Since the surplus in energy has to
be divided between the positron and the neutrino there is a continuous energy spectrum for the
positron up to the maximum energy. The mean energy of the positron is roughly 40% of the
maximum energy. So there is always a finite distance between the place of decay and place of
annihilation. For a radionuclide like 18F, with a low maximum energy of 0.635 MeV, the
maximum positron range in tissue (water) is 2.3 mm. The mean range in water is 0.6 mm. Nearly
all other positron emitting radionuclides have a larger maximum energy.
The other limitation is the change from measuring line-integral projection into measuring the
27
place of annihilation directly by means of a time of flight (TOF) measurement. As stated before,
time differences of 3 ps have to be measured to obtain a spatial resolution of 1 mm. Detectors
with these timing properties and also a high sensitivity for 511 keV gamma radiation are not
known (yet). In solid state physics research has been done and is going on but an appropriate
detector material has not been found up till now. The most recent material is LSO with a decay
time of 40 ns and good other parameters for being used in PET scanners.
Fig. 13.
FDG brain scan with the experimental Siemens/CTI HRRT (MPI Collogne),
LSO/GSO detectors with DOI information. Spatial resolution ~ 2 mm FWHM
(Courtesy Siemens/CTI).
4.
Possibilities of PET in research and patient care
The clinical applications of PET are in the fields of cardiology, neurology and oncology. In the
cardiology the measurement of the myocardial blood flow under rest and stress conditions with
13
N-ammonia and the energy consumption with 18FDG is a standard examination in order to
discriminate between ischemic and infarcted tissue. In the neurology the cerebral blood flow
and/or the energy consumption of the brain is the standard examination. In the oncology PET is
used for the detection of tumors and to measure the effect of therapy on the tumor metabolism.
28
4.1
Applications in cardiology
In table 7 different radiopharmaceuticals for cardiac studies are summarized. 13N-ammonia is
used for the measurement of the myocardial blood flow. To study the viability of the heart it is
used in combination with 18FDG. The combination of ammonia rest, ammonia stress and
metabolism study deliver a much too large number of images to evaluate individually. For this
reason software to re-orient the images perpendicular to the long axis of the heart (left ventricle
in fact) followed by a translation of the data into quantitative parameters of blood flow and
glucose consumption per heart region has been developed. Blood flow and metabolism are than
visualized per examination in a so called polar map. It is also possible to use the electro-cardiac
signals to make a gated cardiac study. From this data it is possible to generate images of the
beating heart and if from these images the wall of the left ventricle can be detected, the wall
motion can be quantified.
Table 7. Radiopharmaceuticals commonly used in cardiology
Measurements
Radiopharmaceutical
Blood flow
Metabolism
Receptor density
Hypoxia
H215O, 13NH3, 82Rb
18
FDG, 11C-fatty acids,
11
C-CGP
18
F-fluoromisonidazol
11
C-acetate
4.2
Applications in neuroscience
In table 8 different radiopharmaceuticals for neuroscience studies are summarized. The clinical
and research programs in Groningen are directed to glucose metabolism (18FDG), protein
synthesis rate (PSR) with 11C-tyrosine and blood flow with H215O. The improvement in resolution
can be seen in fig. 12 where the glucose metabolism of the brain is shown for the different
generations of PET scanners. The progress which still can be made using new detector
technology can be seen when comparing this with the results of the HRRT research scanner as
shown in fig. 13. For oncological studies both 18FDG as well as L-[1-11C]tyrosine are available.
Software for the translation of measured radioactivity into glucose-consumption (18FDG) and
protein synthesis rate of 11C-tyrosine have been developed. The D2 receptor in the human brain
can be studied with 18F-DOPA or 11C-raclopride and is of importance in the case of Parkinson's
disease. Measurement of the regional cerebral blood flow (rCBF) with H215O is of great
importance to discover the functional anatomy in fields like cognitive neuroscience, linguistics,
selective attention and to measure the effect of drugs on the rCBF in different categories of
patients. Stimulus research brings together nearly all fields: from biophysics through psychiatry
to language. Also the merging of functional PET images with the anatomical MRI images, using
segmentation techniques, is of importance to localize the functional anatomy.
29
Table 8. Radiopharmaceuticals commonly used in neuroscience
Measurements
Radiopharmaceutical
Blood flow (rCBF)
Blood volume
Oxygen extraction
Glucose metabolism
Tumor metabolism
Receptor measurements
H215O, C15O2
CO, C15O
Combination of above
18
FDG
18
FDG, 11C-amino acids
11
C-methylspiperon, 18FESP,
raclopride
18
FDG, H215O
Stimulus research
11
18
F-DOPA,
11
C-
4.3
Applications in oncology
In table 9 different radiopharmaceuticals for oncological studies are summarized. For the study of
tumor metabolism 18FDG is often used but other possibilities in the form of amino acids do exist.
Also the effect of therapy on the tumor metabolism can be quantified by measuring before and
after therapy. By performing the second study already during the therapy may be also a
prognostic statement can be made. For oncological brain studies the use of an amino acid can be
favourable due to the better signal to noise ratio which can be obtained with respect to the
glucose metabolism study. It is also possible to generate "whole-body" images by projecting a
number of consecutive transverse images into a planar image. Also the PET/CT combination into
one device is of importance for the combination of functional and anatomical images in
radiotherapy and surgery.
Table 9. Radiopharmaceuticals commonly used in oncology
Measurements
Radiopharmaceutical
Tumor perfusion
Tumor metabolism
H215O, 13NH3
FDG, 11C-tyrosine,
dine, 18FLT
11
C-cytostatics
Cytostatica kinetics
Therapy evaluation
18
11
C-methionine, 11C-thymi-
4.4
Applications outside medicine
The labeling of new pharmaceuticals to establish their distribution in vivo is a field where the
pharmaceutical industry can get new information the distribution in vivo and on the possible
efficacy of new drugs before large clinical studies are started. Large investments can be
prohibited in case of a negative finding or better rationalized in case of a positive finding.
Measuring the effect of new pharmaceuticals by measurement of differences in blood
30
flow/glucose consumption/protein synthesis rate etc. before and after therapy. Dose effect
relationships can be studied also.
In chemical reactors sometimes the exact localization of processes is completely unknown. Using
nuclear techniques and PET special it is possible sometimes, by scaling down reactor structures
to within the opening size of a PET scanner, to obtain this information which can be used in the
optimization process of the reactor.
5.
Data analysis
A wide variety of data analyses is possible depending on the data acquisition protocol. One has to
realize that there are only a few measurable quantities on which the data analysis, often leading to
functional parameters, is dependent. In this part only general remarks are made on some different
possibilities in data analysis. In fact each radiopharmaceutical has its own procedure for a
quantitative data analysis.
5.1
General evaluation of PET images, visual, non-quantitative
The PET scanner is able to deliver quantitative images (Bq/voxel) if corrected for attenuation and
scatter. Evaluation can be done by visual inspection as the most simple analysis. Also ratio's can
be calculated by drawing ROI's (Regions Of Interest) in the images. Ratio's of e.g.
striatum/cerebellum, left/right can be calculated. Also the content or better the concentration of
activity of the ROI (Bq/ml) can be compared to the injected dose (Bq/g). The ratio of these two is
the Differential Absorption Ratio (DAR) or Specific Uptake Value (SUV). If the DAR or SUV
has a value of 1 it means a homogeneous distribution. If something more elaborate e.g. a
functional parameter like blood flow (ml/min/100g), glucose consumption (µmol/min/g) is
wanted a more complex analysis scheme has to be set up.
There are only a few measurable quantities
i)
Radioactivity (after all corrections) in tissue (Bq/voxel). The chemical identity is only
known at the moment of injection, due to metabolism different metabolic species can be
formed. The radioactivity in tissue information is available per voxel per unit of time or
per time frame.
ii)
One can take (arterial) samples and make a biochemical analysis and measure the
radioactivity of the different radioactive species. Radioactivity in plasma per chemical
identity as function of time is than obtained. The content of the arterial plasma with the
original radiopharmaceutical as function of time is also known as the "input function".
The PET-scanner should be "cross calibrated" with the well counter in which the samples
are measured. The short half life makes fast chemical analyses necessary.
31
5.2
General compartimental models
In the biological system we have to do with a steady state system. Regardless of the fact that we
have a continuous flow of materials, or of energy, the concentrations (or the chemical states) in
the different compartments do not vary. The concentrations remain constant even though there is
a flow of material along the chain. Positron emitting radiopharmaceuticals are normally
compounds with a high specific activity (sa). If the radiopharmaceutical is chemically identical to
endogenous substances or is in competition with endogenous substances we have to do with
compartment systems with distribution volumes and constant transfer flows. With X* the amount
of radioactive material is represented and with X the amount of non radioactive material.
X* + X = X,
specific activity = X*/{X* + X} = X*/X
Carrier free: when producing X* no X is produced. A non carrier added (nca) synthesis: during
synthesis of X* no carrier X is added but some is possible included in the starting environment,
e.g. CO2.
Quite a number of different compartment model can be thought of: not at a steady state, open or
closed models depending on radiopharmaceutical used and in which system. In this part only the
case of a system with n strongly connected compartments, 1, 2, 3, ...n, is examined. The material
(x0) is introduced in compartment 1 at t=0. A rapid mixing is assumed and then the material will
distribute itself throughout the various compartments. At time t these will contain X1, X2, .... Xn.
The transfer constant from compartment i to compartment j is symbolized by kij.
The variations of the quantity of material in compartment i as function of time is than given by:
d Xi /dt = Σj kjiXj - Σj kijXi
with kij = 0 when i=j
To solve these first order, linear with constant coefficient, differential equations it is convenient
to Laplace transform these equations.
With xi the Laplace transform of Xi this translates to:
sx1 - x0 - Σj=2 kj1xj + Σj=2 k1jx1 =0
for i=1
sxi - Σj kjixj + Σj kijxi =0
if xi(0) = 0 for i > 1
So, ordinary linear algebra with n equations with n unknowns has to be solved. To solve this the
determinant of the set of equations has to be solved. Then, after the inverse Laplace transform,
the solution are of the form:
X1 /x0 = Σp Ap exp(-apt)
Xi /x0 = A1 + Σp=2 Ap exp(-apt)
The coefficients AP and aP are determined from the roots of the determinant. In this way multicompartment models can be solved. In practice only a limited number of quantities can be
measured, as discussed before, and the number of free parameters and so the number of
32
compartments, has to be kept rather limited in order to obtain a unique solution of the problem.
5.3
The FDG model.
2-[18F]Fluoro-2-deoxy-D-glucose or FDG is the most frequently used PET radiopharmaceutical.
Its behaviour can be described by three compartments: FDG in plasma (CP), FDG in tissue (CE)
and FDG-6P (FDG-6-phosphate) in tissue (CM). After this phosphorilation step the FDG is
recognized as not the right glucose and metabolism stops. This is a complicating factor because at
first the behaviour is identical to the endogenous glucose but after a first metabolic step the
distribution volume changes dramatically. In due course the FDG will de-phosphorilated but this
step is not important in the first hour after injection when the measurement is done. So there are 4
transfer constants, k1 and k2 between plasma and tissue and k3 and k4 between tissue and
metabolized FDG. As said, the distribution volume of FDG will be part of the distribution volume
of glucose and this complicates the quantification of the glucose metabolism of course.
So we get the following differential equations, for simplicity the *, as marker for radioactive part,
has been removed.
In the tissue there is in-flow and out-flow from plasma and metabolite compartment:
dCE(t) /dt = k1CP(t) - (k2 + k3)CE(t) + k4CM(t)
The metabolite compartment is only connected to the tissue compartment:
dCM(t) /dt = k3CE(t) - k4CM(t)
After Laplace transform the linear equations are:
sCE (s) - CE (0) = k1CP (s) - (k2 + k3)CE (s) + k4CM (s)
sCM (s) = k3CE (s) - k4CM (s)
Relating (CE) and (CM) to (CP):
{[s2 + s(k2 + k3 +k4) + k2k4]/ (s + k4)} CE (s) = k1CP(s)
or, after finding the roots of the quadratic expression,
{[(s + a1)(s + a2)]/(s + k4)} CE(s) = k1CP(s)
and {[(s + a1)(s + a2)]/ (k1k3)} CM(s) = CP(s)
The inverse Laplace transform yields:
CE(t) = k1/(a2 - a1) {(k4 - a1) exp(-a1t) + (a2 - k4) exp(-a2t)} O CP(t)
CM(t) = k1k3/(a2 - a1) { exp(-a1t) - exp(-a2t)} O CP(t)
O denotes the operation of convolution:
33
a(t) O b(t) = f a(τ)b(t-τ)dτ
The measurable quantity is the tissue concentration CI with CI(t) = CE(t) + CM(t), the plasma
concentration of FDG and the plasma concentration of the (cold) glucose. As said before, FDG
and glucose are competitive substrates for hexokinase in the phosphorelation process and their
rates follow the Michaelis-Menten relationship. This competition is hidden in the so called
"Lumped Constant" or LC. So based on the three measurable quantities mentioned above the
LCMRGlc (Local Cerebral Metabolic Rate of Glucose consumption) can be evaluated.
LCMRGlc = (CPc/LC) (k1k3/(k2 + k3) with Cpc the "cold" value
The problem is the lumped constant (LC): which value for what tissue has to be taken and what to
do in pathological tissue. Few values have been determined, most frequently the value of 0.42 is
used.
A less complicated method can be followed if (i) the individual rate constants are not of interest
(ii) k4=0 and (iii) a dynamic scan is available. If this is the case a graphical procedure can be
followed: the Gjedde-Patlak method.
k1k3/(k2 + k3) = {CI(t)/CP(t) - k1k2/(k2 + k3)2}/ 0ft CP(t')dt'/CP(t)
A linear relationship between the ratio of tissue and plasma concentration and the integrated
plasma concentration divided by its actual value is established and LCMRGlc can be calculated
from the slope, the cold plasma concentration and the Lumped Constant as given above. Since the
three conditions mentioned above are normally fulfilled the Gjedde-Patlak approach is the the
commonly followed approach.
5.4
Receptor modeling
The aim of receptor studies is not only to visualize the receptor distribution in a 3D volume but
also make an absolute quantification in terms of receptor density Bmax and affinity constant KD.
The ligand used for receptor studies should have a high selectivity and affinity for the receptor
system to be studied and should have a low non-specific binding. In receptor studies, due to the
often low amount of receptors, the specific activity plays often a crucial role. Sometimes the
specific activity should explicitly been taken into account: X* + X > X. Sometimes a unique set of
parameters can be found by using multi-injection protocols with injections with different specific
activities or even "cold" material. These variations can be essential to obtain not only a unique
solution but also a solution with small error bars on the individual parameters.
The model used should give a correct description of the data. An equilibrium approach can be
used if the ligand shows a rapid dissociation and a kinetic model analysis of the dynamic tissue
activity curves is an appropriate approach if there is a slow dissociation of the ligand from the
receptors. The compartment model to be used depends on the properties of ligand used of course.
When the ligand upon entering the brain tissue is bound there immediately and irreversibly only
the transport rate k1 can be determined. If the ligand does not bind in the tissue but is simply
clearing back into the blood a equilibrium between blood and tissue activity will be reach
34
asymptotically. The equilibrium will determined by the ratio of k1/k2. In this model the tissue is
described without any specific receptor binding. If the ligand binds to a receptor a third
compartment with rate constants k3 and k4 has to be added. In case of irreversible binding k4 = 0
and asymptotically the accumulation rate k1.k3 /(k1+k3) will be reached (see also the FDG model)
where k3 represents the product of association rate kon and the density Bmax of the receptors. The
actual magnitudes of the different rate constants determine what actually can be learned about the
receptor binding, e.g. if k3 >> k2 the rate limiting step is k1 and nothing can be learned about the
receptor binding. If a ligand with reversible binding is used an equilibrium between blood and
tissue activity will be reached asymptotically and is given by k1/k2 (1+k3/k4) where k3/k4 =
Bmax/KD or also called the binding potential. Mathematically it is easy to add further
compartments, however, since the measurable quantities are rather limited it is very difficult to
extract more than 3 to 4 model parameters unambiguously. Sometimes it is quite difficult to
establish the parameters with the required accuracy. For that reason also multi-injection protocols
with different specific activities are necessary to obtain the required accuracy. It also possible to
use a graphical approach like Gjedde-Patlak plot for the analysis but no individual rate constants
can be determined from such a graphical analysis. In the approaches as described above always
the plasma concentration is one of the measurable parameters. To avoid this blood sampling
procedure the reference tissue procedure can be followed. This procedure is possible if there is
some well defined brain tissue in the field of view which does not have any of the receptors to be
studied. This region devoid of receptors can be used as an input function for the total region. In
this way there are two compartment models: i) a two compartment model for the reference tissue
and ii) a three compartment model for the receptor containing tissue.
The concentration for the reference tissue is described by:
dCr /dt = kr1Cp - kr2Cr
The free ligand in receptor concentration in the receptor containing tissue is described by:
dCf /dt = k1Cp - k2Cf - k3Cf + k4Cb
The specifically bound concentration is described by:
dCb /dt = k3Cf - k4Cb
Where: Cp = the plasma concentration (metabolite corrected)
Cr = reference tissue concentration, reference tissue does not contain the
receptor of interest
Cf = concentration of free (not specifically bound) ligand
Cb = concentration of specifically bound ligand
This set of equation can be fitted from the measurable quantities.
Different approaches are possible and the choise will be dictated by the specific properties of
ligand to be used and the receptor system to be studied.
5.5
Non-quantitative rCBF studies in stimulus experiments
In stimulus studies an experiment is repeated in one volunteer a number of times (e.g. 2
conditions with 6 vs. 6 studies in each condition) and the same experiment is done in a number of
35
volunteers (e.g. 8 - 12) when subtle stimuli are used (e.g. language experiments). The PET
scanner should be used in the 3D-mode (highest sensitivity). How to analyze the studies and to
identify the functional brain areas involved with a statistical significance? The practical standard
at the moment is the method as developed at the Functional Imaging Laboratory (FIL, Institute of
Neurology, London): the Statistical Parametric Mapping (SPM) software package.
The stimulus experiments should be set up very carefully in order not to vary to many parameters.
Preferably only one parameter should be changed. Often one condition is the rest condition and
the second condition is the stimulus condition. If the rest condition is not a well defined condition
a rather confounding situation can be created.
After data acquisition the following steps are made:
i)
prepare rCBF images into right format with corrections if necessary
ii)
realign images per volunteer, calculate mean image
iii)
normalize images into e.g. Talairach space based on the AC-PC lines
iv)
smooth images (noise reduction)
Now all images of the different volunteers are on the same scale and subtraction images can be
made.
v)
Using a General Linear Model a regression analysis can be made and statistical
significances can be calculated and the areas of functional anatomy can be identified.
An example of a study with also the combination of structural MRI images to identify the
anatomical location is the following study. The language localization in cases of left temperal
lobe arachnoid cysts was investigated to see whether these cyst have led to reorganization of the
language function using measurement of the rCBF. Words or letter strings were presented
centered on a screen for 750 ms each, a fairly slow pace which according to the pretest allowed
subjects to understand relative complex sentences. The screen was 90 cm from the subject eyes.
Four right handed patients and four matched volunteers, each 6 vs. 6 rCBF studies were made and
analysed within the SPM framework in order to find the speech center. T1-MRI studies were
made of all patients and volunteers. The T1-MRI studies were segmented into pseudo rCBF study
and the segmented T1-MRI and PET rCBF study were matched by a correlation algorithm. So we
have PET-rCBF studies with the localization of language identified and matched with the
individual MRI study. As is obvious from fig. 14 there is no evidence for an inter-hemispheric
reorganization.
6.
Shielding and dosimetry
6.1
Radiation shielding of accelerator and laboratory
The radiation sources are the places of beam losses: the deflector and slits inside the cyclotron,
slits in bending magnets and of course the target itself. When the accelerated charged particles hit
material gamma radiation and neutrons are generated. Other particles generated by nuclear
reactions have such short range that they are absorbed in the material in which they are generated.
36
Fig. 14.
The individual statistical results from the PET scans, yellow areas, from the
patients projected on the corresponding normalized (Talairach and Tournoux
stereotactic reference system) anatomical MRI slices. The yellow areas are the
language localization areas as found with statistical significance from the PET
rCBF measurements. No evidence for inter-hemispheric reorganization was
found. (Courtesy Stowe et al, Brain and Language 75(2000)347-358)
37
Cyclotrons with proton energies between 15 MeV < Ep < 50 MeV require about the same
shielding because the energies of gamma-rays and neutrons are the same and amount does not
vary to much: (p,xn) reactions. Typically 6 MeV is available for de-excitation by gamma-rays.
On the average one has to calculate with three gamma-rays of about 2 MeV. The maximum
neutron energy is roughly 11 MeV available based on Q-values and beam energies. Most
neutrons are emitted in a forward cone with a half top angle of 300. The neutrons are thermalized
and than disappear by neutron capture.
Table 10. Radionuclides identified in solid materials irradiated in and around accelerators,
adapted from Patterson and Thomas.
Material
Radionuclides
Plastics, oil
Concrete, aluminum
Iron, steel
7
Copper
Be, 11C
as above, plus 22Na, 24Na, 32P, 42K, 45Ca
as above, plus 44(m)Sc, 46,47,48Sc, 48V, 51Cr, 52(m),54,56Mn, 57,58,60Co,
57
Ni, 55,59Fe
as above, plus 65Ni, 61,64Cu, 63,65Zn
For shielding from neutrons hydrogen is the most effective material. For gamma-rays high
density material with high Z-value is the most effective. So most vaults make then cheapest
compromise and are build of ordinary concrete (2.35 g/cm3) with a wall thickness of 1.5 - 1.8 m.
Sometimes a shielding of barite concrete (extra heavy) or concrete with extra water content (extra
stopping for the neutrons) is used. Also a pure water jacket as shielding around the cyclotron is
used in some cases.
In a recent study on the "Evaluation of the radiological and economic consequences of
decommissioning particle accelerators" it was concluded that, based on measurements of acquired
samples from accelerator vaults, especially 152Eu, and into a less extent 60Co,
154
Eu,134Cs,22Na,54Mn, are present in the concrete of the vault. These radionuclides are all
produced by neutron capture reactions and their concentration is also a function of the depth in
the concrete wall. The choice of material in beam stops, collimators and irradiation set-ups can
greatly influence the number of unwanted secondary neutrons produced and by this the amount of
radionuclides produced. In the case studies mentioned in the report the presence of natural Eu (1
ppm) in the sand and aggregates of the concrete and natural Co in the reinforcement bars (100
ppm) is causing these contaminations. Since only a few case studies are presented in this report
one should take his own decisions based on actual measurements at the time of decommissioning.
In the construction of the vault special attention has to be paid for:
i)
Entrance of vault construction via maze or door
ii)
Feed through of cabling, water cooling, ventilation: construct ducts or labyrinths with at
least three legs (or two bendings of 900)
38
iii)
Air-borne radioactivity and the ventilation of vault
iv)
Cooling water of cyclotron in a closed primary system should be cooled by a secondary
system
Evaluate the need of radioactive waste tanks or only delay tanks for all the water coming
from the radiochemical lab
Evaluate the need for storage of solid radioactive waste
v)
vi)
Radiation safety at the radiochemical PET-laboratory
i)
Always used lead shielding of 5 cm or more
ii)
Calculate thickness hot cells based on expected activity, e.g. 100 GBq, and on allowed
radiation level at 10 cm from the outside of the hot cell wall (1 µSv/h)
iii)
Install radiation monitors
iv)
Install sufficient ventilation
v)
Fulfil the under pressure requirements
vi)
Install the same clock (time) at laboratory and PET camera
6.2
Dosimetry
The International Committee on Radiological Protection (ICRP) gives the international
recommendations on radiation. The ICRP publications give all relevant information. The
radiation dose can be calculated conform the recipe as developed by the Medical Internal
Radiation Dose (MIRD) Committee of the Society of Nuclear Medicine (SNM).
The MIRD system handles the concept of source and target organs. Each organ is a target for
radiation from a radiopharmaceutical deposited in the body. An organ that contains radioactivity
is not onlay a source but also a target due to self-irradiation. The radiation dose received by target
organs depends on the geometry inside the body and the nature and energy of the radiation.
Dm = A0τ S
with Dm = mean dose
A0 = administered activity
τ = residence time = effective time that the administered dose spends in the source organ
S = dose to the target from unit cumulated activity in the source organ (dose per unit
cumulated activity). S values are tabulated for many source organs in combination with
various target organs.
The mean absorbed dose per unit administered activity to the target organ rk is:
Dm(rk) / A0 = Sh th S(rhk)
with S(rhk) the contribution of source organ rh to target organ rk
39
The cumulated acitvity Acumh is the sum of all nuclear transitions in organ h during the time
interval of interest:
Acumh = f Ah(t) dt
If the activity function Ah(t) can be approximated by a sum of exponentials it can be written as:
Ah = Sj Bj exp{-(λ + λj)t}
with l the physical decay constant and lj is the biological decay constant
The effective half-life is defined by Tjeff = ln2/( λ + λj)
The residence time in source organ h is defined as: th = Acumh / A0
where A0 is the administered activity and Acumh is the cumulated activity in organ h. What this
definition of residence time in fact means is that the integrated value of the time activity curve is
used to calculate a time during which administered activity was present organ h.
Using the residence times in the different sources organs one is than able to calculate the doses in
all target organs.
To come a weighted calculation for the whole body first the mean dose equivalent in the target
organ is calculated. The mean dose equivalent HT in a target organ or tissue T is given by: HT =
DT Q N (ICRP 53) with DT the mean absorbed dose, Q the quality factor, N is the product of any
other modifying factors. For electrons and gamma radiatiation the quality factor is one and
modifying factors are normally not an issue. The effective dose equivalent is than calulated by
adding the weighted organ or tissue mean dose equivalents, see table 11:
HE = ST wT HT
Table 11. The weighting factors for calculation HE are:
Tissue
wT
Gonads
Breast
Red bone marrow
Lungs
Thyroid
Bone surfaces
Remainder (5 tissues)
0.25
0.15
0.12
0.12
0.03
0.03
0.30
A convenient way to do these calculations is to use the MIRDOSE 3 personal computer software
as developed by the Radiation Internal Dose Information Center of the Oak Ridge Institute for
Science and Education, Oak Ridge (TN), USA.
In dosimetry still a mixture of old units and the new SI units is in use. The relation of the most
40
frequently used units is given below.
Dosimetry units:
Röntgen
1 R = 2.58E-04 C/kg
rad
D rad = 0.869 X R
1 rad = 0.01 J/kg (= 100 erg/g)
rem
D rem = RBE.X rad
Gray
1 Gy = 1 J/kg = 100 rad
Sievert
D Sv = RBE.X Gy
RBE = Relative Biological Effect = 1 for gamma-rays and electrons
RBE > 1 for neutrons and alpha particles
The cyclotron as available now for PET centers is different from the older machines used in the
field of nuclear physics not only because of the limited number of particles and fixed energy but
also because of the incorporation of the targetry for the most important radionuclides.
Automation and computer control is integrated into the design. Not only the beam quality but also
the beam current is a major parameter because the current determines the production capacity.
Beam quality is not that crucial for radionuclide production and in fact the beam power density
(W/cm2) should not be too high. For the day-to-day operation no separate operating team is
required, the cyclotron can be operated by the technical chemical staff. The present developments
tend into a few directions, but reduction in costs by e.g. a reduction in maximum beam energy is a
general goal for the marketing of cyclotrons for clinical PET-centers. Reduction in maximum
energy results in a smaller machine and consequently a reduction in costs is possible. Lower
energy leads also to less penetrating particles and consequently to thinner targets with a lower
yield. Consequently the target technology becomes more difficult and more critical by this
reduction in beam energy because the beam current has to be increased to keep up in production
capacity. For experimental nuclear physics super conducting cyclotrons have been built. The
change to a superconducting magnet decreases the weight with a large factor. This technology
has also been used for a commercial PET production cyclotron. Since the production capacity
should stay the same the thickness of the shielding is also the same but the overall size off the
vault, and so the costs, can be decreased again. Some cyclotron manufacturers also provide local
movable shielding of concrete and lead, fitting tightly around the accelerator, resulting in lower
total mass of the shielding. Also developments in linear accelerators (linacs) and in Radio
Frequency Quadrupole accelerators (RFQ's) for the production especially of the four PET
radionuclides, are taking place. However, it still has to be shown that this, may be cost effective
solution, is a competitor in radionuclide yields with the now operating 17 MeV proton cyclotrons.
The radiation detectors used in positron cameras at the moment are made of BGO in most cases
but also NaI and BaF2 has been used or still is in use. Although BGO and BaF2 have a high
stopping power for 511 keV and a number of other favourable properties, the light yield of NaI is
also favourable. The ideal detector for a positron camera should have a time resolution of less
than 10 ps and this combined with other properties like high stopping power, high Z,
non-hygroscopic etc.. This extreme fast timing would allow for the measurement of the place of
annihilation within a few millimeters by means of the time-of-flight (TOF) measurement. With
the present detectors only the line on which the annihilation took place is being determined. The
filtered back-projection reconstruction technique in combination with block structure of the
41
detectors makes a spatial resolution of 4-7 mm FWHM standard. Recently LSO (lutetiumorthosilicate) has been discovered as a scintillator. LSO combines the good properties of BGO
with high light yield (75% of the yield of NaI) and is also rather fast (40 ns). A disadvantage is
the presence of natural radioactive isotope of lutetium but, since a coincidence technique is
employed, this will not influence the image formation. The higher light yield will improve the
energy resolution and by this decrease the scatter fraction. At the moment small LSO PETscanners are being built for small animals (rats and mice) and a spatial resolution of 2 mm
FWHM has been achieved in these systems. The first commercial whole body machine is now
also available, see table 4 in 3.6.2, though not yet at this superior resolution. Also GSO, already
used in older scanners, is now used as a continuous detector in a new designed PET scanner, see
table 4 in 3.6.2.
In order to solve the matching problems between PET and CT images combined PET/CT
machines have been developed. In fact these are a separate PET and CT put together in one
housing operated independently but under one general software system so the images have the
same pixel size etc., see table 6 in 3.6.3. In this way the changes in positioning which always will
happen if a patient is transferred from one scanner to another are not there anymore. The
advantages of combining functional and anatomical images are in the fields of head/neck surgery,
pulmonary surgery and also in radiotherapy, especially when conformal therapy is considered.
8.
PET and other imaging modalities
The most common imaging technique in medicine uses X-rays. In its most simple form a density
projection is generated by holding the subject of interest between the X-ray tube and a
photographic plate. The advanced form can be found in a CT-scanner in which a rotating X-ray
source and detectors make a transverse section image. Again sort of a density map is generated
although extraction of the exact density will not be possible, and is also not necessary for
diagnostic use, due to the broad energy spectrum of the generated X-rays. Due to the large
difference in density between bones and tissue the bones can be visualized perfectly while small
differences in tissue density will be more difficult to visualize. The use of contrast agents, like
fluids with high densities and high Z-components, can change the difference in density and by
this the interpretation of the images dramatically.
The Nuclear Magnetic Resonance (NMR) technique is in use to visualize the protons (bound to
water) in the human body. Homogeneous magnetic fields up to 3 T are in use in medical NMR
scanners, nowadays abbreviated to MRI (Magnetic Resonance Imaging). In order to have a short
imaging time, gradient fields with frequency decoding are used. The strength of the NMR signal
is proportional to the difference in population of the spin-up and spin-down state. Under normal
conditions at room temperature the ratio between spin-up and spin-down is rather close to unity.
The NMR technique is a rather insensitive technique for this reason but which is successful
because of the high water concentration in the human body. Also paramagnetic contrast agents
like Gd-DTPA are used to increase the contrast. Both, the X-ray and the MRI technique, supply
anatomical information.
With NMR also information on the structure of molecules can be obtained, as is done in the
chemistry. This also possible in the human body but limited to the brain and to molecular
structures which have a concentration of 0.1 mM or more as a rule of the thumb. The limitation to
mainly brain tissue is because the signals from water and fatty tissues have to be suppressed and
these concentrations are rather low in the brain in contrast to e.g. the thorax.
42
Functional MRI, or fMRI, uses the effect of deoxyhemoglobin on the MRI signal, the so called
blood oxygen level-dependent (BOLD) signal change. This signal is interpreted as changes in
rCBF and is used for stimulus experiments. In this fMRI and PET are sometimes seen as
competitive approaches. However, the mechanisms underlying the BOLD signal change are not
fully understood yet. The quality of fMRI as a research tool can be judged by the results of the
experimental applications and also by a cross-validation with the gold standard: H215O-PET
measurements for measuring the rCBF. Since sensitivity, spatial resolution, temporal resolution
etc. of both techniques is quite different it suggests that a complementary approach will probably
yield the best results.
By using radionuclides bound to different molecular structures functional emission imaging
became available in the 1950's. It is functional imaging because the chemical structure and the
human metabolism determine the faith of the molecule in vivo. PET is the ultimate form of
nuclear medicine. Because the nuclear reactions for the production of positron emitting
radionuclides are of the type (p,n) or (p,α) in most cases, the element produced is different
from the target element. By this type of reactions the amount produced in weight is extremely
low (1 GBq of 11C has a weight of 35 pg) while the amount of radioactivity is considerable
(50-100 GBq of 11C can routinely be made and a patient dose is approximately 400 MBq).
This so called specific activity ( MBq/µg) is of importance e.g. for receptor research and
makes it possible to call it "tracer" experiments.
Since the PET method supplies functional information the combination with X-ray and NMR
techniques, CT and MRI, would yield an identification of the functional anatomy. In order to
make this combination the images of the different disciplines should be available in a
transparant way and image resize and re-orientation techniques should be available to match
the images from the different modalities. At the moment there are no general applicable
routines available to perform this kind of matching operations. In a number of PET centers the
combination of PET-, CT- and NMR-images is subject of interest. In some institutions also
the comparison and transformation of PET images to a stereotactic brain atlas have been
performed. The head and brain are the first structure/organ of choise to evaluate the "multimodality" matching for obvious reasons. Below the brain it very difficult due to differences in
positioning inside the scanner. For that reason now the first commercial developed PET/CT
machines are entering the market and their possibilities will be evaluated in the coming years.
References:
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45