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292
Positron emission tomography: physical principles, instrumentation and performance evaluation
The Slice Thickness Test
The slice thickness test determines the real slice thickness for the system.
The Homogeneity Test
The homogeneity test determines the CT values for all kV settings.
The Noise Test
The noise test determines the pixel noise for all kV settings.
The MTF Test
The MTF test determines the modulation transfer function for the system.
Radiation and sensitivity profile widths
The radiation profile describes the distribution of radiation energy within a continuous
medium along a line parallel to the scanner rotational axis. The radiation energy peaks in
the center of the field and falls off diffusely towards the edges due to scatter and collimator
penumbra. The magnitude of penumbra depends on the z dimension of the focal spot (parallel
to scan rotation axis) and focus collimator- patient geometry. A ready pack film placed
horizontal to the tabletop and at the CT centre was exposed using all available slice thickness.
Exposed films were scanned using a film scanner with 0.01 cm step size and 285 pixel
resolutions. The full width at half maximum (FWHM) of optical density profiles can be
measured for all corresponding slice thickness. These FWHM values represent the radiation
profile widths.
X-ray generator
The x-ray generator is an integral part of the CT system, and in most newer systems is
specifically designed for CT. Typical CT systems operate at relatively few tube potential
settings since each requires different calibration data for image reconstruction. Tube currents
range from 10-600 mA, with mAs values typically between 100-1000 mAs. Tests on the Xray generator include evaluation of peak potential (kVp), timer accuracy (s), mAs linearity
and reproducibility. Linearity of mAs for different kVp can be verified by obtaining the
product of dose and time at different mA and time settings.
Image Quality
A reconstructed CT image is essentially a map of energy weighted x-ray attenuation values
in the scanned slice. Its accuracy in a practical scanner is constrained by intrinsic physical
limitations of the system design. The phantom for measurement of image quality and slicethickness is placed in position inside the CT gantry is shown in figure 12 on the left. The
uniformity of CT slice image for water is on the right side in figure 12. Slices may be
generated at various levels of the phantom to check the image quality.
Positron emission tomography: physical principles, instrumentation and performance evaluation
Figure 12: Slice-thickness phantom in the CT gantry and the uniform CT image.
293
Finite sampling of the image space imposes limits on object spatial frequencies reproduced
in the image. Dose constraints, and limits on x-ray tube output and detector efficiency, cause
statistical uncertainties in attenuation measurements. Superimposed on these are random
and systematic errors from a variety of sources.
Radiation measurement
In a CT scanner, rotation of the x-ray source results in a band of radiation extending
partially or completely around the patient. Within the slice volume, the dose distribution
exhibits symmetry similar to that of arc therapy. The shape of the distribution is a function
of scan arc, which varies from 180° to >400° in clinical systems. Hence it is essential to
determine the radiation dose to tissues under different CT scan conditions. This can be done
by placing CT head dosimetry phantom on the patient table and align phantom axis parallel
to scan rotational axis with slice plane through mid length of phantom. Standard head scan
with a slice width of 5-6 mm should be taken and in the image all three holes in alignment
rod should be clearly seen. Tape the phantom in position when aligned and the ion chamber
can be placed in a hole near the surface (1 cm depth hole) corresponding to maximum dose
point. If rotation angle of scan mode is <360°, then rotate the phantom so that one dosimeter
position (at 1 cm depth) is aligned with midpoint of scan arc. If scan angle is >360°
(overscan) then position the dosimeter at middle of overscan region. If this position is
unknown, take several measurements with chamber at different positions. Note that for
scanners with continuous rotation capability, the maximum dose position varies from one
scan to the next; consult manufacturer before proceeding. Measure exposure for all standard
head techniques, at all calibrated kVp settings, and with any operator selectable beam filter
combinations (appropriate for head scanning). Also measure for all available slice widths.
294
Positron emission tomography: physical principles, instrumentation and performance evaluation
The multiple scan average dose (MSAD) values should be determined for all standard scan
conditions, and for any variable, which might be expected to influence the dose.
In interpreting the data, determine whether the change in parameters produce the expected
result on dose, e.g., a linear change with mAs or slice width. It should be ensured that the
measurement conditions are as per specifications of the manufacturer or comparable with
them.
References
1. Phelps ME. PET: the merging of biology and imaging into molecular imaging. J Nucl Med 2000;
41:661–81.
2. Juweid ME, Cheson BD. Positron emission tomography and assessment of cancer therapy. N Engl J
Med 2006; 354: 496-507.
3. Kluetz PG, Meltzer CC, Villemagne VL, Kinahan PE, Chander S, Martinelli MA, et al. Combined
PET/CT Imaging in Oncology. Impact on Patient Management. Clin Positron Imaging 2000; 3:223-
230.
4. Sweet WH. The use of nuclear disintegration in the diagnosis and treatment of brain tumor. N Engl J
Med 1951; 245:875-878.
5. Brownell GL, Sweet WH. Localization of brain tumors with positron emitters. Nucleonics 1953;
11:40-45.
6. Ter-Pogossian MM, Powers WE. The use of radioactive oxygen-15 in the determination of oxygen
content in malignant neoplasms. In: Radioisotopes in Scientific Research. Pergamon Press.London.
1958.
7. Nutt R. The history of positron emission tomography. Mol Imaging Biol 2002; 4:11-26.
8. Berko S, Hereford FL. Experimental studies of positron interactions in solids and liquids. Rev Modern
Phys 1956; 28:299- 307.
9. Harpen MD. Positronium: Review of symmetry, conserved quantities and decay for the radiological
physicist. Med Phys 2004; 31: 57-61.
10. Hoffman EJ, Huang SC, Phelps ME, Kuhl DE. Quantitation in positron emission computed tomography:
4. Effect of accidental coincidences. J Comput Assist Tomogr 1981; 5:391-400.
11. Muehllehner G, Karp JS, Surti S. Design considerations for PET scanners. Q J Nucl Med 2002;
46:16-23.
12. Humm JL, Rozenfeld A, Del Guerra A. From PET detectors to PET scanners. Eur J Nucl Med Mol
Imaging 2003; 30:1574-1594.
13. Casey ME, Nutt R. A multi-slice two-dimensional BGO detector system for PET. IEEE Trans Nucl
Sci 1986; 33:760-763.
14. Colsher JG. Fully three-dimensional positron emission tomography. Phys Med Biol 1980; 25:103–
115.
15. Phelps ME, Cherry SR. The changing design of positron imaging systems. Clin Positron Imaging
1998; 1:31-45.
16. Melcher CL, Schweitzer JS. Cerium-doped lutetium oxyorthosilicate: a fast, efficient, new scintillator.
IEEE Trans Nucl Sci 1992; 39:502-505.
17. Nutt R. For: Is LSO the future of PET? Eur J Nucl Med Mol Imaging 2002; 29:1523-1525.
Positron emission tomography: physical principles, instrumentation and performance evaluation
295
18. Karp JS. Against: Is LSO the future of PET? Eur J Nucl Med Mol Imaging 2002; 29:1525-1528.
19. Yang Y, Dokhale PA, Silverman RW, Shah KS, McClish MA, Farrell R, et al. Depth of interaction
resolution measurements of a high resolution PET detector using position sensitive avalanche
photodiodes. Phys Med Biol 2006; 51:2131-2142.
20. Herschman HR. Micro-PET imaging and small animal models of disease. Curr Opin Immunol 2003;
15:378-384.
21. Bloomfield PM, Rajeswaran S, Spinks TJ, Hume SP, Myers R, Ashworth S, et al. The design and
physical characteristics of a small animal positron emission tomography. Phys Med Biol 1995; 40:1105–
1126.
22. Chatziioannou AF. PET scanners dedicated to molecular imaging of small animal models. Mol Imaging
Biol 2002; 4:47-63.
23. Tai C, Chatziioannou A, Siegel S, Young J, Newport D, Goble RN, et al. Performance evaluation of
the microPET P4: a PET system dedicated to animal imaging. Phys Med Biol 2001; 46:1845-1862.
24. Jeavons A, Chandler RA, Dettmar CAR. A 3D HIDACPET camera with sub-millimetre resolution for
imaging small animals. IEEE Trans Nucl Sci 1999; 46:468-473.
25. Kemp BJ, Kim C, Williams JJ, Ganin A, Lowe VJ. NEMA NU 2-2001 performance measurements of
an LYSO-based PET/CT system in 2D and 3D acquisition modes. J Nucl Med 2006; 47:1960-7.
26. Derenzo SE. Mathematical removal of positron range blurring in high-resolution tomography. IEEE
Trans Nucl Sci 1986; 33:565-569.
27. Zanzonico P. Positron Emission Tomography: A Review of Basic Principles, Scanner Design and
Performance, and Current Systems. Semin Nucl Med 2004; 34: 87-111.
28. Cherry SR. The 2006 Henry N. Wagner Lecture: Of mice and men (and positrons) advances in PET
imaging technology. J Nucl Med 2006; 47:1735-45.
29. Budinger TF. PET instrumentation: what are the limits? Semin Nucl Med 1998; 28:247-267.
30. Muehllehner G, Karp JS. Positron emission tomography. Phys Med Biol 2006; 51:R117-37.
31. Daube-Witherspoon ME, Muehllehner G. Treatment of axial data in three-dimensional PET. J Nucl
Med 1987; 28:1717-1724.
32. Hudson HM, Larkin RS: Accelerated image reconstruction using ordered subsets of projection data.
IEEE Trans Med Imaging 1994; 13:601-609.
33. Daube-Witherspoon ME, Matej S, Karp JS, Lewitt, RM. Application of the 3D row action maximum
likelihood algorithm to clinical PET imaging. IEEE Trans Nucl Sci 2001; 48:24-30.
34. Miller TR, Wallis JW: Fast maximum-likelihood reconstruction. J Nucl Med 1992; 33:1710-171.
35. Motta A, Damiani C, Del Guerra A, Di Domenico G, Zavattini G. Use of a fast EM algorithm for 3D
image reconstruction with the YAP-PET tomography. Comput Med Imaging Graph 2002; 26:293-302.
36. Schoder H, Erdi YE, Larson SM, Yeung HW. PET/CT: a new imaging technology in nuclear medicine.
Eur J Nucl Med Mol Imaging 2003; 30:1419-1437.
37. Townsend DW, Beyer T, Blodgett TM. PET/CT scanners: a hardware approach to image fusion.
Semin Nucl Med 2003; 33:193–204.
38. Beyer T, Townsend DW, Blodgett TM. Dual-modality PET/CT tomography for clinical oncology. Q J
Nucl Med 2002; 46:24-34.
39. Conti PS, Lilien DL, Hawley K, Keppler J, Grafton ST, Bading JR. PET and [18F]-FDG in oncology:
a clinical update. Nucl Med Biol 1996; 23:717-735.
40. Hoh CK, Schiepers C, Seltzer MA, Gambhir SS, Silverman DH, Czernin J, et al. PET in oncology:
Will it replace the other modalities? Semin Nucl Med 1997; 27:94-106.
296
Positron emission tomography: physical principles, instrumentation and performance evaluation
41. Ling CC, Humm J, Larson S, Amols H, Fuks Z, Leibel S, Koutcher JA. Towards multidimensional
radiotherapy (MD-CRT): biological imaging and biological conformality. Int J Radiat Oncol Biol
Phys 2000; 47:551- 560.
42. Blodgett TM, Meltzer CC, Townsend DW. PET/CT: form and function. Radiology. 2007; 242: 360-
85.
43. Jeavons AP. Small-animal PET cameras. J Nucl Med 2000; 41:1442-1443.
44. Tornai MP, Jaszczak RJ, Turkington TG, et al. Smallanimal PET: advent of a new era of PET
research. J Nucl Med 1999; 40:1176-1179.
45. Chatziioannou AF, Cherry SR, Shao Y, Silverman RW, Meadors K, Farquhar TH, et al. Performance
evaluation of microPET: a high-resolution lutetium oxyorthosilicate PET scanner for animal imaging.
J Nucl Med 1999; 40:1164-1175.
46. Anger H O. Survey of radioisotope cameras. J Nucl Med 1966; 5: 311–34.
47. Brownell G L, Burnham C A, Wilensky S, Arnonow S, Kaemi H, Strieder D. New developments in
positron scintigraphy and the application of cyclotron-produced positron emitter. IAEA Proceedings
of a Symposium on Medical Radioisotope Scintigraphy, Salzburg, Austria, 1968.
48. Muehllehner G, Buchin MP, Dudek J H. Performance parameters of a positron imaging camera. IEEE
Trans Nucl Sci 1976; 23: 528–37.
49. Surti S, Karp J S, Muehllehner G, Raby P S. Investigation of Lanthanum Scintillators for 3-D PET.
IEEE Trans Nucl Sci 2003; 50: 348–54.
50. Moses WW. Time of flight in PET revisited. IEEE Trans Nucl Sci 2003; 50: 1325–1330.
51. Conti M, Bendriem B, Casey M, Chen M, Kehren F, Michel C, et al. First experimental results of
time-of-flight reconstruction on an LSO PET scanner. Phys Med Biol 2005; 50: 4507–26.
52. Moses WW, Ullisch M. Factors influencing timing resolution in a commercial LSO PET camera.
IEEE Trans Nucl Sci 2006; 53: 78–85.
53. Karp J S, Kuhn A, Perkins A E, Surti S, Werner M E, Daube-Witherspoon M E, et al. Characterization
of a time-of-flight PET scanner based on Lanthanum Bromide. IEEE Nucl Sci Symp Conf Record
2006; 1919–23.
54. Pichler BJ, Judenhofer MS, Catana C, Walton JH, Kneilling M, Nutt RE, et al. Performance test of an
LSO-APD detector in a 7-T MRI scanner for simultaneous PET/MRI. J Nucl Med. 2006; 47:639–647.
55. Watanabe M, Shimizu K, Omura T, Sato N, Takahashi M, Kosugi T, et al. A high-throughput whole-
body PET scanner using flat panel PS-PMTs. IEEE Trans Nucl Sci 2004; 51:796–800.
56. Zaidi H, Mawlawi O, Orton CG. Point/counterpoint. Simultaneous PET/MR will replace PET/CT as
the molecular multimodality imaging platform of choice. Med Phys 2007; 34:1525-1528.
57. Handler WB, Gilbert KM, Peng H, Chronik BA. Simulation of scattering and attenuation of 511 keV
photons in a combined PET/field-cycled MRI system. Phys Med Biol 2006; 51:2479–2491.
58. National Electrical Manufacturers Association. NEMA Standards Publication NU-2 1994: Performance
Measurements of Positron Emission Tomographs. Washington, DC: National Electrical Manufacturers
Association; 1994.
59. Daube-Witherspoon ME, Karp JS, Casey ME, et al. PET performance measurements using the NEMA
NU 2-2001 standard. J Nucl Med. 2002;43:1398 – 1409.
60. Performance Measurements of Positron Emission Tomographs. Rosslyn, VA: National Electrical
Manufacturers Association; 2001. NEMA Standards Publication NU 2-2001.
61. American Association of Physicists in Medicine (AAPM) Report No. 39, Specification and acceptance
testing of computed tomography scanners. American Institute of Physics, New York, 1993.
PET Radiochemistry and PET
Radiopharmaceuticals
M.G.R. Rajan
PET radionuclides
A large majority of the radionuclides used in industry and medicine are produced either in
the nuclear reactor or in a cyclotron where the nuclear reactions take place transmuting one
element to another. Several radionuclides have an established role in nuclear medicine
procedures both for diagnosis and therapy. The latest molecular imaging with PET requires
the use of radiopharmaceuticals (RPs) labeled with positron emitting radioisotopes (PER).
The PET imaging at the molecular level bridges the gap between laboratory science and
clinical medicine by providing the most specific and sensitive means for imaging molecular
pathways and interactions in human tissues. Due to its relevance in several diseases,
particularly cancer, PET-imaging is poised to make a major scientific as well as financial
contribution to drug development, particularly for neurological diseases and cancer treatment.
PET imaging requires the use of ‘true metabolites’ i.e., sugars, amino acids, fatty acids etc.,
essentially made of H, C, N and O which the cells in the body can metabolize. The conventional
131
radiopharmaceuticals (RPs) labeled with
purpose. One looks for the radioactive isotopes of H, C, N and O, for imaging, which are
produced in a cyclotron. The short-lived positron emitters such as 11C, 13N and 15O are cyclotron
produced. A suitable radioisotope of H is not available but, fortunately, 18F (also cyclotron
produced) can substitute H in several metabolites.
I,
99m
Tc, 67Ga,
111
In etc., are not suitable for this
Apart from 18F, the workhorse of PET-RPs, only a few other short-lived radionuclides are
useful for PET imaging. Their production is described here along with a few others expected to
be clinically useful in the future. The physical properties of these isotopes are given in table 1.
29 7
PET Radiochemistry and PET Radiopharmaceuticals298
Table 1: Properties of commonly used and promising PET isotopes and method of their production
Nuclide Physical Mode of Photon Abundance Common
half-life decay (%) energy (%) production method
(T
) (keV)
1/2
11
C 20.4min + (100) 511 200
18
F 110min + (97)EC(3) 511 194
68
Ga (From
68
Ge generator) 68min + (89)EC(11) 511 178
124
I 4.2 d + (23)EC(77) 511 46
14
N(p,)11C
10
B(d.n)11C
18
O(p,n)18F
68
Zn(p,n)68Ge
124
Te(p,n)
124
I
603 61
1691 10.4
13
N 10min + (100) 511 200
15
O 2min + (100) 511 200
64
Cu 12.7 h + (17.9) 511 35.8
– (39) 578 39
EC(43.1) 1345 0.47
12
C(d,n)13N
16
O(p,)13N
14
N(d.n)15O
15
N(p,n)15O
64
Ni(p,n)64Cu
64
Ni(d,2n)64Cu
68
Zn(p,n)64Cu
1666 43.2
82
Rb(From 75s + (95) 511 190
82
Sr generator) EC(5) 777 13.4 (spallation reaction)
94m
Tc 52min + (70) 511 140
98
Mo 82Sr
94
Mo(p,n)
94m
Tc
EC(30) 871 94
1521 4.5
1868 5.7
Important PET Radionuclides and their production
Carbon-11
Carbon-11 (T½ = 20 min) is produced by the 14N(p,)11C reactions in the cyclotron using
(>10MeV) protons, but may also be produced by deuterons with 10B(d,n)11C, and 11B(p,n)11C
reaction with B2O3 as the target. However 14N is the preferred target as it is easily available. The
14
N2 gas is mixed with a small amount of oxygen (0.5%) so that 11CO2 is predominantly produced
which is separated from the target gases by condensing it in a liquid nitrogen trap and later
removed by warming and flushing with helium. Curie quantities need to be prepared since a
major portion of it decays during synthesis because of the short physical half-life. The 11CO2 is
used as a precursor for compounds, such as 11C-palmitate used for myocardial PET-imaging.
PET Radiochemistry and PET Radiopharmaceuticals 299
11
CO2 is converted to 11C-methoxide and then reacted with HI (hydroiodic acid) to form 11Cmethyl iodide. The latter is the common route for the preparation of most 11C-labeled PET-RPs.
Since this process is cumbersome it is recommended to use 14N2 mixed with traces of H2 in the
target so that 11CH4 is produced directly during irradiation and is easily iodinated to 11CH3I, by a
gas phase iodination with heated iodine-vapours. The 11CH4 can also be converted to 11C-HCN by
heating it with ammonia over a platinum catalyst and used for the production of aliphatic amines
and amino nitriles to be used as intermediates for a number of 11C-PET-RPs.
Copper-64
Copper-64 (T½ = 12.7 d) is receiving much attention as it can be coupled to a variety of
molecules, peptides and proteins by chelation. It decays by a combination of -, + and electron
capture (EC). The decay by EC mode gives rise to a significant amount of low energy Auger
electrons. Thus, 64Cu is considered to be a potential isotope for PET imaging as well as for
therapy. Further the biochemistry of copper in the body is well understood. It is fairly easy to
prepare 64Cu with enriched 64Ni target by irradiation with (>9 MeV) H+ beam, with nuclear
reaction 64Ni(p,n)64Cu and isolated from the target chemically. Since 64Ni can be recovered with
95% yield, the production of 64Cu becomes inexpensive.
The production of 64Cu by the bombardment of natural Zn with protons is also described and
used commercially. However both 64Cu and 67Cu are produced in the process in a ratio of 75:25
with reactions 68Zn(p, n)64Cu and 68Zn(p,2p)67Cu. Purification from the other isotopes formed
during irradiation of other isotopes of Zn present in natural Zn is necessary. The neutron
induced 64Zn(n,p)64Cu in a reactor is also described but has low cross section and is not the
preferred method for producing 64Cu.
Fluorine-18
The most common method of 18F (T
= 110 minutes) production is by irradiating H
1/2
target with high energy protons (9 to 18 MeV) to facilitate the nuclear reaction 18O (p,n) 18F.
Isotopically enriched
18
O-water (usually 95 – 97%) is the target material. Several commercial
manufacturers supply more than 95% enriched 18O-water. A metal target holder (e.g. silver,
titanium, nickel, copper, stainless steel or niobium) is used to hold the H
18
O in a cavity of 0.3 to
2
3ml volume depending on the target design. Following irradiation, the 18O-water along with 18Ffluoride is passed through a bicarbonate anion-exchange resin column to trap the 18F-fluoride.
The unused 18O-water is collected for purification and reuse. 18F- ion is eluted out of the column
with potassium carbonate solution or tertiary butyl ammonium bicarbonate and used for further
synthesis.
The 18O-water should have minimum ionic and non-ionic impurities as some of them,
particularly; the halogens can react with and corrode the silver target body at the high temperatures
and pressures reached during irradiation. The no-carrier added (NCA) specific activity of
18
F-fluoride is very high ~ 74000GBq/g (2000Ci/g), i.e., 74 GBq (2 Ci) of activity is obtained
18
O
2
PET Radiochemistry and PET Radiopharmaceuticals300
from ~1 ng of 18F. In comparison the stable 19F- present in commercial 18O-water can be a 1000
times higher; upto 2g/ml (2ppm). Apart from corroding the expensive silver target bodies over
a period of time, the 19F- can also compete with the 18F for the precursors during synthesis of 18Flabeled compounds. While this can be taken care of by using excess precursor during synthesis,
as for 18F-FDG, but may not be possible where high specific activity of the PET-RP is important,
(e.g., receptor imaging). The suppliers are now willing to produce 18O-water with <0.1 mg/L of
fluoride if specifically asked for.
The 18F-fluoride is mostly used for labeling deoxyglucose to produce 18F-2-fluorodeoxyglucose
(18F-FDG). Other 18F- labeled pharmaceuticals of clinical utility are discussed later.
Germanium-68 (Generator for Gallium-68)
Germanium-68 (t
= 270.8 days) is a cyclotron produced isotope, which could be used as a
1/2
sealed transmission source in PET scanners for attenuation correction prior to PET imaging. It
can be produced only in high energy cyclotron (>40 MeV) by the 66Zn(,2n)68Ge reaction, or the
spallation reactions in a molybdenum target with high energy protons. It decays by electron
capture and remains in equilibrium with 68Ga (t
= 68 minutes). The 68Ge-68Ga generator is
1/2
routinely used to produce PET radiopharmaceuticals for imaging patients with neuroendocrine
tumors in many centers.
Iodine-124
Iodine-124 (T
is produced by the irradiation of solid target made of 96% enriched
protons or deutrons with nuclear reactions
124
I by + is low (23%) while decay by EC is 77%. Photons from the latter mode with energies
>511 keV cause some loss of resolution in PET imaging.
of 4.2 d) has promise since it can be used to make iodinated PET-tracers. It
1/2
124
Te(d,2n)
124
I or
124
Te(p,n)
124
Tellurium oxide with
124
I. However, the decay of
Nitrogen-13
Nitrogen-13 (T½ = 10 min) has been found useful as simple molecules like 13NH3 or other
13
N-labeled molecules requiring very short synthesis time. It is easily produced by irradiating
16
H
O with protons to produce 13N with 16O(p,)13N nuclear reaction. Other method wherein 13C
2
(as gaseous compound or as a solid metal-carbide) is irradiated with H+ is also reported but is
neither convenient nor easy. Proton bombardment of H
nitrites, ammonia, and hydroxylamine. All are reduced to 13NH3 using a combination of Devarda’s
alloy and concentrated sodium hydroxide solution. Methods have been reported wherein 16Owater containing a small amount of ethanol is irradiated with H+ to directly give 13NH3. The
method works by the quenching of free radicals (produced during irradiation) by ethanol,
preventing the formation of nitrates and nitrites, but allows the nascent 13N to combine with H
free radicals to give 13NH3.
In both the methods, the liberated
13
NH3 is dissolved in saline, sterile filtered and used. It
16
O gives mostly nitrates along with
2
PET Radiochemistry and PET Radiopharmaceuticals 301
forms NH
+
ion in aqueous solution and is used in PET imaging for myocardial perfusion.
4
Glutamine and asparagine have been reported to be labeled with 13NH3 for assessment of tissue
viability.
Oxygen-15
The very short-lived Oxygen-15 (T
(with 1% 16O2 gas) with (8-10MeV) D+ particles to give 14N(d,n)15O reaction. With cyclotrons
that can only produce H+ beam, 15O can be produced by the 15N(p,n)15O reaction. 15O2 is mixed
with H2 and heated over a catalyst to give 15O-water useful for cerebral and myocardial perfusion
studies.
=2 min) is produced by irradiating gaseous nitrogen
1/2
Strontium-82 (Generator for Rubidium-82)
Strontium-82 (T
82
Sr82Rb generator. 82Sr is produced by the 85Rb(p,4n)82Sr reaction using the high-energy H
beam or by directing a very high energy H+ beam on 98Mo target to form 82Sr along with a
spallation of neutrons. Since the reaction requires very high energy protons, a typical PET
cyclotron is not suitable.
About 100mCi (3.7 GBq) of the 82Sr (purified from the target) is loaded on a stannic oxide
column in the generator. The 82Rb is eluted as chloride from the 82Sr’82Rb generator with 0.9%
sodium chloride solution. The major radionuclidic impurities in the 82Sr sample are 85Sr (T
65d) and 83Sr (T
significantly less when the 85Rb target is used. 83Sr is a not a worry since it decays to a negligible
quantity by the time 82Sr generator is produced or use.
= 25.6 d) is a convenient parent for Rubidium-82 (T
1/2
= 32 hours). The major contaminant is 85Sr when 98Mo target is used, but is
1/2
75 sec), in a
1/2 =
1/2
=
+
Technetium-94m
Technetium-94m (t
= 52 minutes) is produced by the nuclear reaction, 94Mo(p,n)
1/2
using an enriched 94Mo target in a medical cyclotron. It is rapidly separated from 94Mo by steam
distillation. Being a + emitter,
94m
Tc is receiving attention as it may be a substitute for
in SPECT radiopharmaceuticals and used for PET-imaging with improved sensitivity and
resolution.
94m
99m
Tc
Tc
Synthesis of PET Radiopharmaceuticals
PET radiopharmaceuticals (RPs), have radionuclides that are positron emitters, by virtue of
which they differ from the conventional
medicine. The PET-isotopes: 11C, 15O, 13N and 18F used for labeling ‘true metabolites’ are shortlived and this places several constraints on the synthesis time for the PET-RPs, quality control
and their clinical use. The unique advantage of PET is that it often depicts a true representation
of biological processes after in vivo administration. For example 18F-fluorodeoxyglucose (FDG)
is an analog of glucose used for cellular metabolism and H
99m
Tc-labeled and other SPECT-RPs used in nuclear
15
O for cerebral perfusion.
2
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