Apparatus for the diagnosis of body structures into which a gammaemitting radioactive isotope has been introduced

Body structures into which a gamma-emitting isotope has been introduced are investigated using an imaging system comprising in sequence a collimator, a first detector, a filter, and a second detector. The first detector is partially transparent to the emerging photons and the filter is adapted to have an absorption edge energy fractionally less than the energy of the radioactive isotope source. The first detector produces an image of both scattered and non-scattered photons and the second detector simultaneously produces an image of only the scattered photons. The two images can be manipulated to produce an image due only to non-scattered photons which is the image of interest.

The present invention relates to apparatus for the diagnosis of body 
structures into which a gamma-emitting radioactive isotope has been 
introduced. 
Gamma detectors and associated collimators are used in nuclear medicine to 
study the morphological and/or physiological changes in biological 
organisms. The best possible spatial resolution is required to study 
objects within an organism, such as a cancer tumor, as small as a few 
millimeters in size. 
One of the serious limitations to the spatial resolution of detectors in 
current use is due to the scattering of photons on the electrons of the 
organism, known as Compton scattering. Scattered photons are deflected and 
may be misinterpreted as if they were emitted in another part of the body. 
This distortion leads to decrease of image contrast and to the creation of 
artifacts, which are details or structures present in the image but having 
no counterpart in the organism. Therefore, for accurate diagnosis, the 
scattered photons should be eliminated when the image is reconstructed. 
The cross-section for Compton scattering decreases with energy and 
therefore only those radioisotopes with energies above 70 keV are usually 
used. 
When scattered, photons loose part of their energy. In order to reject the 
scattered photons during the image reconstruction, it is essential to use 
quasi-monochromatic sources of gamma radiation and photon detectors with 
very good energy resolution. While the radioactive gamma sources are 
essentially monochromatic, the energy resolution of most presently 
available photon detectors is poor; for example, NaI scintillators have an 
energy resolution of 15-20%. Only the semiconductor detectors have an 
energy resolution of better than 10%, but they have the disadvantage of 
being much more expensive. 
In my French patent application No. 77 01150 published as Pat. No. 23 77 
642 I have proposed a new nuclear medical imaging arrangement which allows 
a significant improvement in energy resolution (.DELTA.E/E.ltoreq.5%). 
This improvement is obtained by the use of detectors comprised of elements 
of high atomic number (but thinner than those hitherto proposed, with the 
high atomic number element's being selected to have an absorption edge 
energy slightly below the radiation energy of the gamma-emitting 
radioisotope which is to be detected. 
In my French patent I have also proposed a way of improving the images 
obtained with existing detectors used in nuclear medical imaging, such as 
the NaI scintillation camera. This is achieved by the use of filters of 
high atomic number materials in conjunction with energymatched 
gamma-emitting isotopes. This enables an image to be formed substantially 
only of photons which are emitted from the isotope within the tissue and 
scattered by the tissue. A further image can also be formed in the normal 
way without the filter and this further image can be considered to be a 
composite of the undesired false image due to photons scattered within the 
tissue, superimposed on the real image representing the distribution of 
the photon emitters inside the tissue. A mathematical procedure permits 
the image due to scattered photons to be subtracted from the composite 
image, thereby obtaining the real image. 
However, with this technique two images have to be formed at different 
times which is disadvantageous (possibility of movement of the patient so 
that the two images are not truly coincident period of time required 
etc.). 
The principal object underlying the present invention is to provide an 
improved nuclear medical imaging arrangement and method which avoids the 
need to form a plurality of images at different times, which yields 
improved images and facilitates interpretation of these images. 
In order to satisfy this object there is provided, in accordance with the 
present invention, apparatus for the diagnosis of body structures into 
which a gamma-emitting radioactive isotope has been introduced, the 
apparatus comprising, in sequence, a collimator through which photons 
emanating from the body structures being studied are directed; a first 
photon detector, said first photon detector being adapted to absorp a 
fraction of the photons incident thereon; a filter comprising an element 
with an atomic number not less then 74; a second photon detector for 
detecting photons passing through said filter and analysing means for 
producing an image due to non-scattered photons from the image of 
scattered photons formed by the second detector and the image of scattered 
and nonscattered photons formed by the first detector. 
Thus two detectors are used, with a heavy element filter interposed between 
them, so that two images--one that of the scattered photons alone and the 
other a composite of the real and false images--may be obtained 
simultaneously. The true image of the photons emanating from within the 
tissue is obtained from these two images (for example, by selecting upper 
and lower scintillator detectors of differing pulse decay times and 
utilizing suitable electronic circuitry). The arrangement results in good 
energy resolution and the images are obtained in a much more efficient 
manner. 
Further advantageous arrangements are set out in the accompanying subclaims 
.

Table 1 shows the absorption edge energies of certain elements with atomic 
number Z greater or equal to 74. The following quantities are of interest 
and are listed in Table 1: 
.mu..sub.+ --absorption coefficient just above E.sub.a, 
.mu..sub.- --absorption coefficient below E.sub.a, 
.mu..sub.15% --absorption coefficient at E=0.85.times.E.sub.a. 
Table 2 is a listing of radioactive isotopes tabulated against the suitable 
very high atomic number element from which the filter should be built. 
Before describing the specific arrangement of the present invention it is 
considered helpful to discuss the underlying theory and the arrangements 
proposed in my French patent. 
It is well known that the mass absorption coefficient of an element 
increases drastically at some characteristic energy (see FIG. 2 for lead). 
The absorption edge energy is between a few keV for low atomic number 
elements and as high as 115 keV for Uranium, i.e. it increases for higher 
atomic numbers. In medical applications, photons of low energy have only a 
relatively small chance of escaping from the organism and only the highest 
Z elements can be considered as radiation detectors or filters. The 
absorption edge energies for some elements with Z.gtoreq.74 are listed in 
Table 1. Elements like Re, Os, Ir, or the actinides different than Uranium 
are omitted because they are prohibitively expensive. 
The possibility of improving the energy resolution of existing photon 
detectors results from the existence of gamma emitters with energies 
slightly above the absorption edge energy of corresponding detector 
materials (see Table 2). Only isotopes with a decay life longer than 10 
minutes are listed and then only the energy, E.sub.gamma, of radiation 
close to the absorption edge of interest is included, even if these 
radioisotopes are not monochromatic. 
Table 2 shows the decay time (half life time t.sub.1/2), the isotope energy 
E.sub.gamma, and the relative interval between E.sub.gamma and the 
absorption edge E.sub.a. The last column of table 2, lists the element 
whose absorption edge is closest to the energy of the radioisotope. This 
element with atomic number Z is ideal for use as heavy detector or filter. 
However, also elements with Z-1 and Z-2 can be used to reject scattered 
photons. 
FIG. 1 shows one arrangement proposed in my French patent, the layout of 
the apparatus does not differ from the classical arrangement. 
The collimator is transparent to photons propagating in the chosen 
direction only. It makes it possible to obtain a two-dimensional (2-D) 
projection of the 3-D object under study. The collimator is usually 
realised as a "multihole collimator", i.e. a thick sheet of metal in which 
holes have been made. The collimator is a device well known in the prior 
art and will therefore not be discussed in the following. For an arbitrary 
detector, or if the source has an energy considerably higher than the 
absorption edge energy, both the "true", i.e. unscattered and the "false", 
i.e. the scattered photon will be absorbed inside the detector. Actually, 
the probability of absorbing a false photon is higher than that of 
absorbing a "true" photon. For example, for the most popular radioisotope 
Tc.sup.99m for which E.sub.gamma =140 keV and a detector consisting of 
pure lead, E.sub.a =88 keV, we have .mu.(E=140 keV)=2.27 g/cm.sup.2, 
whereas for photons scattered through an angle of 45.degree. we have 
u(E=127 keV), =3.84 gm/cm.sup.2. The differences in absorption coefficient 
is generally inadequate for rejection of the scattered photons. 
Drastically different is the situation for the very particular cases when 
the energy of an emitted photon is very close to the absorption edge of a 
detector of appropriate thickness (as proposed in my French patent). 
Almost all scattered photons will have energies below the absorption edge 
and the mass absorption will be very small. For the sake of simplicity, 
let us assume that the filter consists of a pure,very high atomic number 
element. With an appropriate thickness of the filter, almost all, say 90% 
of "true" photons will be stopped, but almost all scattered photons will 
pass through the filter. This difference will be lost if the detector is 
too thick, that is to say that the filter should be thinner than 5 
gm/cm.sup.2. 
FIG. 3 shows the alternative arrangement proposed in my French patent. Here 
a high atomic number filter Z.gtoreq.74, matched to the emission energy of 
the isotope is removably interposed between the collimator and the 
detector. A first image F due mainly to scattered photons is obtained with 
the filter in position. The filter is removed and a second image NF is 
obtained, this time of the scattered and non-scattered photons. The true 
desired image of the non-scattered photons is then obtained by 
mathematical techniques. The first image F is, in effect, used to correct 
the second image NF. 
Let us assume that in the image obtained without filter the fraction of 
scattered photons is .beta. (usually 20-50%). To account for the scattered 
photons, the filter should have an appropriate thickness. If the filter is 
too thin, some of the unscattered photons will pass through it. This can 
lead to the creation of artifacts when the image is reconstructed. For a 
filter of thickness x, the transmissivities for scattered and unscattered 
photons are t.sub.1 =exp-(.mu..sub.- x) and t.sub.2 =exp (.mu..sub.+ x), 
respectively. It can be shown that in an image obtained with the filter, 
the fraction b of non-scattered photons is 
EQU b=(1-a)/((t.sub.1 /t.sub.2)*a+1) 
where a is the average fraction of photons scattered inside the body. The 
fraction b should be small, say 
EQU b.ltoreq.b.sub.o and 
EQU x.gtoreq.1n((1-a)*(1-b.sub.o)/(a*b.sub.o))/(.mu..sub.+ -.mu..sub.-). 
For example, with b.sub.o =10%, a=10% and the thickness of a lead filter 
should be x.gtoreq.0.48 g/cm.sup.2 .apprxeq.0.42 mm. On the other hand, if 
the filter is too thick, a portion of the scattered photons will be 
stopped. Thus the time necessary to obtain the correction matrix will be 
longer, or the irradiation dose must be higher. 
In the following capital letters will be used for the matrix 
representations of images. Furthermore, the asterisk "*" is used when the 
matrix is multiplied by a scalar or other matrix. The following matrices 
are of interest: 
I--real distribution of radioisotope inside studied biological object; 
NS--image of the object when all scattered photons are rejected ("signal"); 
S--image of the object when only scattered photons are accounted for 
("noise"); 
NF--image obtained without the filter; 
F--image obtained with filter; 
R--reconstructed image. 
The matrix A can be defined in the following way: 
NF=NS+S 
F=t.sub.1 *S+t.sub.2 *NS 
NS=(1-A)*I 
S=A*I 
where t.sub.1 .apprxeq. exp -(.mu..sub.- x) and t.sub.2 .apprxeq. exp 
-(.mu..sub.+ x) are the transmissivities of the filter for scattered and 
unscattered photons, respectively. However, tests have shown that good 
reconstruction can be obtained when the matrix A is replaced by the scalar 
fraction a. For a given radioisotope, a can be estimated from a knowledge 
of the anatomy of the studied person. For radioisotopes with E.sub.gamma 
=80-120 keV, a=10-40% for all organs with exception of the thyroid. Thus 
the optimum reconstruction is: I=(F-t.sub.1 *NF)/((t.sub.2 -t.sub.1) 
(1-a)).sub.2. For example, for a Pb filter t.sub.1 =52% and 27% for x=0.5 
and 1 g/cm.sup.2 respectively. Thus the filter thickness should be between 
0.25 g/cm.sup.2 and 1.0 g/cm.sup.2. 
Filters can be built of pure, very high atomic number elements or their 
compounds. In the following, the term "equivalent thickness of the filter" 
is used, which gives the thickness of only the very high Z elements in 
units g/cm.sup.2 inside the filter. As the other components of the filter 
are supposed to be of low Z their absorption coefficients are negligible 
even if their contribution to the weight is considerable. 
In practice the arrangement of FIG. 1 is scmething of a compromise and the 
arrangement of FIG. 2 results in the need to form two images at different 
times with the aforementioned disadvantages, particularly movement of the 
patient during data acquisition. 
The present invention is, as already indicated directed to an arrangement 
which overcomes this difficulty. The arrangement of the invention makes 
use of two detectors separated by an appropriate filter as shown in FIG. 
4. 
The top detector, closer to the patient, should be transparent, i.e. only 
10-50% of the available photons should be stopped therein. More 
specifically, it is proposed to use a Xe multiwire proportional chamber as 
the top detector, and the Anger camera (Na I scintillator) as the second 
detector. The Xe detectors have very good intrinsic spatial resolution, 
&lt;1mm, and a reasonable energy resolution, .DELTA.E/E.apprxeq.20-25% at 100 
keV. However, they are partially transparent for photons with E.sub.gamma 
.gtoreq.50 keV. This normally severely limits the use of gas detectors in 
nuclear medicine. In the disclosed embodiment of FIG. 4, however, the top 
detector should be transparent, and thus it is very convenient to use Xe 
as the top detector. 
Another class of photon detectors with excellent spatial resolution and 
reasonable energy resolution are detectors based on the liquid noble 
gases, e.g. liquid argon and liquid xenon. Thus the top detector of FIG. 4 
may also be a thin, liquid noble gas detector. 
The FIG. 5 embodiment uses a fast scintillator, such as CsI, BaF, CsF or 
plastic scintillator, as the top detector and an appropriate filter to 
improve the imaging capabilities of the existing Anger camera. I.e. two 
different scintillators, separated by an appropriate filter, are used. 
With this arrangement a transparent filter, e.g. a heavy metal loaded glass 
or plastic is necessary because the use of an opaque filter would mean 
that both the top and bottom scintillator would have to be provided with a 
separate set of photomultipliers which would drastically impair the 
imaging capability. The use of a transparent filter brings the special 
benefit that only one set of photo-multipliers is required. Of course it 
is necessary for the single set of photomultipliers to be able to 
distinguish between photons stopped in the top and bottom scintillators 
respectively. This can be done conveniently by using a fast scintillator 
such as cesium iodide as the top detector and a slow scintillator such as 
NaI as the bottom detector. The inverse situation is also possible but not 
so convenient because of differences in stopping power. 
By way of example the pulse delay time of CsI=60 nsec. whereas NaI=240 
nsec., and circuits exist which can analyze the time of pulse decay. 
Thus,it is possible to determine whether the gamma was stopped in the top 
(fast) scintillator or in the NaI crystal below the filter. This technique 
is well known--it is called phoswitch--but has never been used for 
position sensitive detectors. Furthermore, the use of an optically 
transparent filter based on very high atomic number elements and placed 
between the two elements of a phoswitch detector is entirely novel. In 
existing phoswitch detectors, the thickness of both fast and slow crystals 
is arbitrary. It should be noted that to practise this invention, the 
thickness of both the top detector and the filter must be chosen in 
accordance with this disclosure, whereas the thickness of the bottom 
detector is not critical. 
The functions of the electronic modules in FIG. 5 can be realised using 
prior art devices. The signals from the photomultipliers are first 
amplified. 
The signals from the amplifiers go to pulse height analysers, such as 
analogue to digital converters and are then passed to position 
localisation circuits which are essentially the same as in existing Anger 
cameras and which identify the positions at which the photons are 
absorbed. This information is passed to the computer. 
In addition the output of the pulse height analysers is passed to the 
energy selection modules which analyse the intensity of the individual 
scintillators and thus the energy of the absorbed photons. This 
information is also passed to the computer. It will be appreciated that 
the operation of the position localisation circuits and the energy 
selection circuits depends crucially on the scintillator in which 
scintillation occurred,this information is provided by time of pulse decay 
modules. 
The time of pulse decay can be realised using a time-to-digital converter 
or, alternatively using constant delay and coincidence circuits. However, 
it should be pointed out that light emission from the fast scintillator is 
smaller than from the NaI crystal. Thus, if the energy selection circuit 
is the same for both detectors, the photons absorbed in fast scintillator 
may be misinterpreted as the scattered photons detected in NaI. Thus a 
crucial element of the present invention is the presence of the 
interconnections between the pulse decay modules and the position 
localisation and energy selection modules. 
Furthermore all the information is stored in the computer, which may be a 
microprocessor, for subsequent off-line analysis. 
The computer also produces the resultant real image from the information it 
receives, i.e. using the matrix analysis previously described. 
TABLE 1 
______________________________________ 
Z E.sub.a u.sub.+ u.sub.- 
u.sub.15% 
______________________________________ 
W 74 69.51 11.4 2.55 3.7 
Pt 78 78.38 9.62 2.09 3.2 
Au x 79 80.67 8.53 1.42 2.5 
Hg x 80 83.08 8.23 1.38 2.4 
Tl 81 85.52 7.93 1.34 2.3 
Pb 82 85.52 7.63 1.30 2.15 
Bi x 83 90.54 7.33 1.26 2.1 
U 92 115.0 4.79 0.865 
1.3 
______________________________________ 
E.sub.a in keV; u.sub.+, u.sub.-, u.sub.15% in g/cm.sup.2 
X u.sub.+, u.sub.-, u.sub.15% calculated by extrapolation. 
TABLE 2 
______________________________________ 
t.sub.1/2 E.sub..gamma. keV 
(E.sub..gamma. - E.sub.a)/E.sub.a 
______________________________________ 
Ir.sup.193 
11 d 69.5 ? W 
Sm.sup.153 
47.1 h 69.8 0.4% W 
Gd.sup.153 
236 d 69.8 0.4% W 
Co.sup.61 
99 m 70 0.7% W 
Cu.sup.61 
3.3 h 70 0.7% W 
Ge.sup.66 
2.5 h 70 0.7% W 
Ba.sup.133 
7.2 y 70 0.7% W 
Pm.sup.151 
27.5 h 70 0.7% W 
Lu.sup.177 
6.8 d 71.6 2.9% W 
Os.sup.185 
93.6 d 71.6 2.9% W 
Pm.sup.145 
18 y 72 3.5% W 
W.sup.187 
24.0 d 72 3.5% W 
Ge.sup.77 
11.3 h 73 4.8% W 
Ho.sup.164 
36.7 m 73 4.8% W 
U.sup.239 
23.5 m 73.6 5.6% W 
Au.sup.193 
15.8 h 73.7 5.7% W 
Sr.sup.83 
33 h 74 6.0% W 
Gd.sup.161 
3.6 m 78 6.8% W 
Bi.sup.204 
11.6 h 78.5 0.15% Pt 
Lu.sup.173 
1.4 y 78.8 0.53% Pt 
Ce.sup.144 
285 d 79.9 1.9% Pt 
Tm.sup.168 
85 d 79.9 2.0% Pt 
Mo.sup.101 
14.6 m 80 2.1% Pt 
I.sup.131 
8.1 d 80 2.1% Pt 
Pr.sup. 145 
22 m 80 2.1% Pt 
Eu.sup.147 
24 d 80 2.1% Pt 
Ho.sup.166 
27.3 h 80 2.1% Pt 
Ir.sup.193m 
11.9 d 80.2 2.3% Pt 
Pd.sup.100 
4.0 d 80.7 2.9% Pt 
Pd.sup.100 
4.0 d 80.7 ? Au 
Se.sup.75 
127 d 80.8 0.11% Au 
Ea.sup.223 
11.3 d 80.9 0.23% Au 
Ee.sup.133 
5.27 d 81 0.33% Au 
Ba.sup.133 
7.2 y 81 0.33% Au 
La.sup.138 
10.sup.11 y 81 0.33% Au 
Te.sup.121i 
140. d 81.9 1.45% Au 
Pt.sup.191 
3.0 d 82.5 2.16% Au 
Po.sup.206 
8.8 d 82.9 2.64% Au 
Pb.sup.211 
36. m 83. 2.75% Au 
Dy.sup.157 
8.2 h 83.1 ? Hg 
Gd.sup.153 
230 d 83.3 0.23% Hg 
Cd.sup.104 
59.2 m 83.5 0.47% Hg 
Kr.sup.79 
1.44 d 84.0 1.1% Hg 
Tn.sup.231 
10.7 d 84.1 1.2% Hg 
Tm.sup.170 
129. d 84.23 1.33% Hg 
Tc.sup.170 
127. d 84.3 1.4% Hg 
Ra.sup.224 
3.64 d 84.3 1.4% Hg 
Tn.sup.228 
1.91 y 84.4 1.5% Hg 
Ta.sup.182 
111. d 84.667 1.8% Hg 
Re.sup.182 
3.0 h 84.67 1.85% Hg 
Re.sup.183 
125. d 84.7 1.9% Hg 
Ta.sup.183 
5.2 d 84.7 1.9% Hg 
Ac.sup. 225 
10. d 85. 2.2% Hg 
Pb.sup.81i 
31.5 m 85 2.2% Hg 
Nd.sup.151 
12. m 85.4 2.7% Hg 
Br.sup.77 
2.4 d 86. 0.56% Tl 
As.sup.77 
38.9 h 86. 0.56% Tl 
Te.sup.160 
72.3 d 86. ? Tl 
Eu.sup.155 
1.8 y 86.4 0.99% Tl 
Ho.sup.161i 
56.3 h 86.4 0.99% Tl 
Pa.sup.233 
27.2 d 86.66 1.25% Tl 
Tb.sup.160 
73.1 d 86.7 1.36% Tl 
Th.sup.233 
22.1 m 86.9 1.59% Tl 
La.sup.142 
77. m 87. 1.7% Tl 
Pd.sup.109 
13.6 h 87. 1.7% Tl 
Cd.sup.109 
470. d 87.6 2.43% Tl 
Er.sup.171 
36.7 m 88.0 ? Pb 
Pb.sup.214 
26.8 m 88.4 0.45% Pb 
Te.sup.123i 
112. d 88.63 0.71% Pb 
Te.sup.127i 
103. d 88.67 0.76% Pb 
Eu.sup.156 
14.7 d 88.9 1.0% Pb 
Lu.sup.176 
2.4 .times. 10.sup.10 y 
88.9 1.0% Pb 
Sb.sup.120i 
5.9 d 89. 1.1% Pb 
Eu.sup.145 
5.0 d 89. 1.1% Pb 
Tb.sup.156 
5.16 d 89.1 1.2% Pb 
Hf.sup.175 
7.0 d 89.4 1.52% Pb 
Ge.sup.69 
1.63 d 90.0 2.2% Pb 
Dy.sup.155 
10. h 90.28 2.3% Pb 
Ho.sup.164i 
36.7 m 90.6 0.087% Bi 
Lu.sup.171 
1.64 y 90.6 0.087% Bi 
Lu.sup.172 
6.7 d 90.6 0.087% Bi 
Bi.sup.204 
12.0 h 90.9 0.42% Bi 
At.sup.209 
5.5 h 90.95 0.47% Bi 
Nd.sup.147 
11.4 d 91.26 0.81% Bi 
Ba.sup.131 
11.5 d 92.0 1.6% Bi 
Ta.sup.182 
118. d 92.0 1.6% Bi 
Er.sup.169i 
105. d 92.1 1.7% Bi 
Th.sup.234 
24.3 d 92.13 1.7% Bi 
Cu.sup.67 
2.44 d 92.2 1.8% Bi 
Ga.sup.67 
3.25 d 92.3 1.9% Bi 
Kr.sup.76 
9.7 h 93.0 2.7% Bi 
Ag.sup.111 
7.54 d 93.0 2.7% Bi 
Ta.sup.182 
111. d 93.0 2.7% Bi 
Hf.sup.181i 
5.4 h 93.3 3.0% Bi 
Cd.sup.107 
6.74 d 93.5 3.2% Bi 
Yb.sup.169i 
30.9 d 93.64 3.3% Bi 
Dy.sup.165 
2.3 h 94.79 4.5% Bi 
Cr.sup.48 
23. h 116.0 0.35% U 
Tl.sup.200 
26.5 h 116.5 0.78% U 
Pb.sup.198 
2.3 h 116.9 1.1% U 
Mn.sup.56 
2.6 h 117. 1.2% U 
U.sup.234 
2.5 .times. 10.sup.5 y 
117.5 1.6% U 
Ga.sup.65i 
15.2 m 118. 2.0% U 
Yb.sup.167 
18.5 m 118. 2.0% U 
Yb.sup.177 
1.88 h 119. 2.9% U 
______________________________________