Embed Size (px)
Citation preview
BMME 560 Medical Imaging: X-ray, CT, and Nuclear Methods
Nuclear Medicine Systems: Basics and Isotopes
SPECT Instrumentation
Guest Lecturer: Marijana Ivanovic
Office: Radiology, 2112 Old Clinic
Tel: 843-0717
Today• Emission vs. Transmission imaging
• Basic (“desired”) properties of radionuclides used for imaging
• Detectors - configuration and characteristics
• Planar Imaging– list mode, static, dynamic, gated, whole body
• SPECT Instrumentation
• Assignment 6
Emission vs. Transmission Imaging
X-ray methods : Transmission Imaging
• Radiation position (direction) is known• Intensity of source is known (known flux (mAs)
and energy (kVp))
X-ray tube
Detector
N0
N=N0e-x
Measure attenuation coefficient ANATOMY
Detector
Emission vs. Transmission Imaging
Nuclear Medicine methods : Emission Imaging Measure concentration and distribution of radiopharmaceutical in the body PHYSIOLOGY (Organ Function, not structure)
• Radiation position (direction) is NOT known• Intensity of source is NOT known• Energy is known
Nuclear Medicine Imaging – Basic Step:
•Production of radionuclide •Labeling of radionuclide with pharmaceutical (tracer)•Injection (or inhalation) of radiopharmaceutical into patient •Wait for distribution and uptake of tracer in the organ of interest
•Imaging (detect -from radionuclide decay)
Tracer Injection Uptake
Time: sec-day
Distribution
Modes of Radioactive Decay:
• Gamma-ray emission (g) - Isomeric Transition (IT) - Internal Conversion (IC)
• Alpha (a) emission • Beta minus (ß-) emission and (ß-, g)• Positron (ß+) emission and (ß+, g)• Electron capture (EC) and (EC, g)
Requirements for Radiotraces
Need -ray emitters (exception:+ emitters – ß+rapidly annihilated with electrons and produce -ray).Charged particles (,-) cannot penetrate tissue for emission imaging.
Requirements for Radiotraces
Desire isotopes relatively simple decay scheme – ideally one or two -rays, no - or -rays. (- or -rays only increase radiation dose to the patient)
Decay scheme complexity:
4399mTc
4399 Tc
(6.01 h)
(1.2 • 105 y)
(140.5 keV - 89 %) Not so good imaging radiotracer
Very good imaging radiotracer
Decay scheme complexity:
Requirements for Radiotraces
If the energy is “too low” a majority of the photons will be attenuated and will not reach the detector simultaneously reduces the signal and adds radiation dose to the patient.
If the energy is “too high” a majority of the photons will pass through the detector without interacting in the detector (also difficult to collimate).
Energy of -rays:
101088664422
2020
00
4040
6060
00
8080
100100
cm of water (muscle)cm of water (muscle)
% t
ran
smit
ted
% t
ran
smit
ted
(511 keV)(511 keV)
99m99mTc (140 keV)Tc (140 keV)
201201Tl (80 keV)Tl (80 keV)
-ray energy should be high enough not to attenuate too much in the body, but low enough to be absorbed by the detector.
Energies of 70-511 keV are used
Requirements for Radiotraces
If the half-life is “too short” it does not permit production, preparation (labeling) delivery, administration and internal distribution for imaging.
If the half-life is “too long” it will take to long to create an image and patient motion will be a problem. Radiation dose to the patient is increases with the half-life, due to a large number of radioactive atoms for a given activity .
Half-life:
Typical Half-lives are on the order of minutes to a few days.
Radiation decays exponentially and its characterized by a “half-life” T1/2 :
A(t) = A0 e- (ln 2* t /T 1/2)
Radiotracer preparation time
Long T1/2
Short T1/2
Radiation dose to the patient ~ with the area under the curve
Imaging Time
Physical Half-life (Tp1/2) and Photon Energy for Radionuclides
commonly used in Nuclear Medicine
Radionuclide Tp1/2 E (keV) Positron E(keV)
Technetium-99m (99mTc) 6.02 hr 140Iodine-123 (123I) 13.3 hr 159Indium-111 (111In) 2.82 d 173, 247Thallium-201 (201Tl) 3.08 d 70, 167Gallium-67 (67Ga) 3.25 d 92, 184, 296Xenon-133 (133Xe) 5.31 d 81Iodine-131 (131I) 8.05 d 364Iodine-125 (125I) 60.2 d 35, 27Fluorine-18 (18F) 109.8 m 511 202Carbon-11 (11C) 20.3 m 511 326Nitrogen-13 (13N) 10.0 m 511 432Oxygen-15 (15O) 2.1 m 511 696Rubidium-82 (82Rb) 1.3 m 511 3150
In order to be useful the radionuclide must be “safe” and able to “trace” within the body, either by itself or attached to a compound.
Requirements for Radiotraces
Must be able to incorporate isotope into a pharmaceutical or other organic compound.
Chemical properties:
R
C
N
R
C
N
R
C
N
R
Tc
C
N
N
R
C
C
N
R
CH2 C O CH3
CH3
CH3
1+
R =
99mTc-Sestamibi
2-[18F] Fluoro-2-Deoxy-D-Glucose
(FDG)
The pharmaceutical part of the radiopharmacetical determines the biodistribution and organ uptake and clearance.
Bone Imaging99mTc-MDP
Cardiac imaging99mTc-Sestamibi
Nuclear Medicine Imaging:
Tumor imaging18F-FDG
Depending on the radiopharmaceutical (radiotracer), different physiological or biochemical functions are being imaged.
Half–times : Physical, Biological, EffectiveThe amount of activity present in an organ after the injection generally changes with time, owing to physical decay of radionuclide and biological uptake and excretion processes.
M = A e-b t = (A0 e-p t) e-b t = A0 e-(b+ p) t = A0
e-eff t
T1/2 eff = T1/2b T1/2 p
T1/2 b + T1/2p
T1/2 eff ≤ shorter of the two, T1/2p and T1/2b
when T1/2p >> T1/2b, then T1/2eff ≈ T1/2b
(tracer excretes very fast – example: 99mTc DTPA, 133Xe)
when T1/2b >>T1/2p, then T1/2eff ≈ T1/2p
(tracer does not excrete or excretes very slowly – examples: 201Tl, 99mTc MAA)Imaging parameters and the amount of activity that could be injected depend on the effective half-life.
Basic Radiation Detector System for Nuclear Medicine Imaging
What do we need to know about the radiation?• Energy?• Position?• How much?
Detector SignalSignal Processing(energy, position..)
Stored to disk
Incoming -ray
What are the important properties of the detector?• Energy resolution• Spatial resolution• Sensitivity• Counting rate
According to the type of information produced:
1. Counters - indicate the number of interactions that occur in the detector
2. Spectrometers – yield the information about the energy distribution of the incident radiation
3. Dosimeters – indicate the net amount of energy deposited in the detector by multiple interactions
Types of Detectors
Detector for NM imaging have to be Counters and Spectrometers.
Types of Detectors
1. Scintillation detectors Inorganic scintillators (NaI(Tl))
Scintillators are materials that emit visible or UV light following the ionization or excitation.
2. Solid State (semiconductor) detectors
Most of the detectors on clinical NM imaging systems are inorganic scintillators.
There are only few cameras with solid state detectors (small field of view),
Properties of Some Scintillator Materials
*Time required for emission of ~67% of the light** Average number of scintillation photons emitted per keV of ionizing radiation energy absorbed.‡ Fast component; # Slow component
1.561.91.822.151.481.85Index of refraction
LittleNoNoNoVeryYesHygroscopic
2250‡
3100#
43004200480039004150 of max. emission ( Å)
2.06.4304.82.540Photo yield** (per keV)
0.856403002.5230Scint. decay time (msec)*
605966745350Effective Atomic No.
-0.6740.8330.955-0.34Attenuation coefficient
(@ 511 keV, cm-1)
4.896.717.47.134.613.67 (g/cm3)
BaF2GSO3LSO2BGO1CsFNaI(Tl)Property
1Bi4Ge3O12 ; 2Lu2SiO5 ; 3Ge2SiO5
Photo Multiplier Tube (PMT)
HighVoltage
Output signal
Inputwindow
PhotocathodeLight Photon
400 V
300 V
500 V. .
. . .
. . .
1200 V
CC
1
3
9
27
81
60000
. . .
e
Focusinggrid
dynode
anode
PMTs are electronic tubes that produce a pulse of electrical current when stimulated by weak light signal.
Scintillation detector consists of a scintillator and a light detector (photomultiplier tube, PMT).
Total electron multiplication is very large : ~ 610 ( ~6x107) for 10 stage dynode.
Inorganic Scintillation Detectors (Scintillator+PMT)
Visible Light
Photons
metal shield
Aluminum orSteel shield
Input window
Glasswindow
-photo
Scintillation center
MgO ili Al2O3
reflectorNaI(Tl)
PMT
NaI - in pure state is scintillator at liquid nitrogen temperaturesNaI(Tl) - scintillator at room temperatures
NaI(Tl) crystal and PMT assemblies
Detectors on majority clinical NM planar and SPECT imaging systems use NaI(Tl) crystals.
Advantages of NaI(Tl) detectors:1. It is relatively dense (=3.67 g/ cm3) and contains an element of
relatively high atomic number (iodine, Z=53). Therefore it is a good absorber and efficient detector of penetrating radiations, such as x-rays and -rays.
2. It is a relatively efficient scintillator, yielding one visible light photon per approximately 30 eV of radiation energy absorbed.
3. It is transparent to its own scintillation emissions. Therefore there is little loss of scintillation light caused by self-absorption, even in NaI(Tl) crystal of relatively large size.
4. A NaI(Tl) detector provides an output signal (from PM tube) that is proportional in amplitude to the amount of radiation energy absorbed in the crystal. Therefore it can be used for energy selective counting.
Detectors on majority clinical NM planar and SPECT imaging systems use NaI(Tl) crystals. NaI(Tl) is excellent for single-photon detectors.
Disadvantages of NaI(Tl) detectors:
1. The NaI(Tl) crystal is quite fragile and easily fractured by mechanical or thermal stresses (e.g., rapid temperature changes). Fractures in the crystal do not necessarily destroy its usefulness as a detector, but they create opacifications within the crystal that reduce the amount of scintillation light reaching the photocathode.
2. Sodium iodide is hygroscopic. Exposure to moisture or a humid atmosphere causes a yellowish surface discoloration that again impairs transmission to the PMT. Thus hermetic sealing is required.
3. Sodium iodide crystals of large size (30-50 cm diam) are difficult to grow and quite expensive.
Semiconductor detectors
Solid-state analogs of gas-filled ionization chambers. When ionizing radiation interacts with the detector, electrons in crystal are raised to
an exited state, permitting an electrical current to flow. 2000-5000 times more dense the gas --> much better stopping power and more
efficient for x- and - rays ( one ionization per 3 eV). Usually requires very high purity materials or introduction of “compensating”
impurities that donate electrons to fill electron traps caused by other impurities. Count individual events Size of electrical signal is proportional to the energy absorbed
Advantages: superb energy resolution
Disadvantages: - high "noise current at room temperature - have to be cooled at 77° K (-196°C)
Limited crystal size ( 5x5 cm) and very expensive
“New” semiconductor detectors
• CdTe and CZT are less well-developed semiconductor materials that overcome two of the major disadvantages of Si and Ge: they can be operated on room temperatures without excessive
electronic noise their high atomic number means that a relatively thin detectors
have a good stopping power for detecting rays.
• Although CdTe and CZT are now being used in some nuclear imaging counting and imaging devices, their use has been restricted to smaller detectors because of difficulty and expense of growing large CdTe and CZT with required purity.
NaI(Tl)PMTPMT
Lead Shield
Lead Shield
F
E
AB
CD
Source
Interactions of photons with a spectrometer:
A - Photoelectric
B - Compton+Photoelectric
C - Compton
D - Photoelectric with characteristic x-ray escape
E - Compton scattering in the shielding and scattered photons enters the detector
F - Characteristic x-ray from lead shield
Sample Spectra with Cs-137
0 200 800Energy (keV)
400 600
32 keV
662 keV
Nu
mb
er o
f in
tera
ctio
ns
10% electron conversionFollowed by a ~32 keV K-shell x-ray
90% 662 keV-ray
0 200 800Energy (keV)
400 600
32 keV
662 keV
Nu
mb
er o
f in
tera
ctio
ns
Due to different (partial deposition) of the energy.Statistical fluctuations in the process by which the energy is deposited and converted into an
electrical signal (random variations in a fraction of deposited energy converted to light, fraction of light that reaches PMT and number of electrons ejected from photocathode per unit energy deposite by the light,…)
Actual energy spectrum Spectrum obtained with a NaI(Tl) detector
Scintillation Camera (Gamma camera)Scintillation Camera (Gamma camera)
Collimator
NaI Crystal
PMT
Lead Shield
Source
Electronic boards
Acquisition &
processing computer
collimatorcollimator
lead shieldlead shield
Gamma camera
CollimatorsCollimators1. To obtain image with the gamma camera, it is necessary to project -rays from the source
distribution onto the camera detector.
2. Gamma rays can not be focused, therefore most practical way to project -rays on an imaging system employs the principle of absorptive collimation for image formation. An absorptive collimation projects an image of the source distribution onto the detector by allowing only those g-rays traveling along certain direction to reach the detector. -rays not traveling in proper direction are absorbed by the collimator before they reach the detector.
3. “Projection by absorption” technique is very inefficient method for utilizing radiation because most of the potentially useful radiation traveling toward the detector is actually stooped by the collimator.
No Collimator
With Collimator
Source
Image of the source
Collimator system is the heart of imaging system – it has the biggest impact on SNR.
Its function is to form an image by determining the direction along which gamma-ray propagates.
L
d
L/2
d d/2
L
Collimator’s Resolution and Sensitivity are determined by the ratio of a collimator hole diameter (d) and length (L).
L (d=const. ) Resolution & Sensitivity d (L=const. ) Resolution & Sensitivity
Density = 3.67 g/cm3
Attenuation coefficient @ 140 keV = 2.64 cm-1 3/8” stops 92% photons
PE fraction = 80%
Scint. decay time = 230 nsec
Photon yield - 40/keV 40*140 = 5,600 light photons emitted for each detected photon
NaI(Tl) Crystal
3/8”(up to 1”) thick
Thinner the crystal better resolution, but lower efficiency.
Resolution & Efficiency vs. Crystal Thickness
Crystal Thickness(Inches)
FWHM(mm)
Photopeak efficiency @140 keV
Photopeak efficiency
@ 511 keV
1/4 3.0 0.70 - 3/8 3.5 0.80 0.055 1/2 3.7 0.85 0.07 5/8 3.9 0.90 0.09 3/4 4.4 0.96 0.10 1.0 4.5 0.99 0.30
SMC
SMC
SM
C SM
C
X- X+
Y-
Y+
PMT’s
NaI(Tl)crystal
Z
XXkX
-+ -=
Z
YYkY
-+ -=
The summing matrix circuits (SMC) combine the signals from the individual PM tubes in such a way that the relative
amplitudes of the X+ and X- signals, and of the Y+ and Y- signals, are proportional to the distance of the scintillation
event from the center line of the crystal. These four signals, are used to determine the location at which the
scintillation event occurred, and the Z signal can be used to determine the energy. Z signal (combined output of from
all PM tubes) is proportional in amplitude to the total amount of light produced by a scintillation event in the crystal.
Spatial Positioning
volt
age
Example of light distribution overPMT locations for 37 tube camera.The scintillations are centeredover the highlighted locations.
0
1
1
0
424
78 74
82 74
82 78
9 11 6
9 9
6
11
9 9
3 3 1
2
2
1
3
3
1 1 1
1
1
1
1
1
1
FOV
1 1 0 0
2 3 3 1 0
3 6 11 5 3 1
2 12 73 67 11 3 1
8 73 423 73 6 2
13 76 73 12 3
6 13 8 2
Energy Windows• Balance between accepting all good events and rejecting scattered photons.• Most camera can acquire multiple (4-8) energy windows simultaneously• Energy resolution of new generation of gamma-cameras is 8-10%.
Gamma Camera Energy Spectra(Summed signal from all PMTs)
Scattered photons are mis-positioned in the image (reduce image contrast)
0
20
40
60
100
80
0 50 100 150 200Energy (keV)
Pu
lse
hei
ght
(a.u
.) Source in air
Source in water
Scattered photons
Image Acquisition
Frame Mode acquisitions:
Static• single or mutiple images acquired at different times and/or different angles• can have multiple energy windows
Dynamic
• series of images acquired sequentially
Gated• repetitive, dynamic imaging (used for cardiac imaging)
Whole Body• Continuous or “step&shoot” table motion during acquisition
Static Acquisition
APAP PAPA
Different views of the same organDifferent views of the same organ
Dynamic acquisition
20 min20 min
1 frame /min1 frame /min
Time
1919171711 33 55 77 99 1111 1313 1515
A
B
C
D
Derdic Dragutin, Hipic, D, 09/02/92
image number
0
1000
2000
3000
4000
5000
0 5 10 15 20
Time
ROI counts
Time
Bladder
R. Kidney
L. Kidney
aorta
Time
Dynamic Acquisition - Processing
Whole Body Bone Scan
Gated AcquisitionGated Acquisition
Time
RR
TTPP
RR
TTPP
RR
TTPP
ECGECG
Ejection Fraction
11 1616
11 1616
Each Image is contribution from 600 heart cycles
Left Ventricle Curve
Projections
ReconstructedTransaxial Slices
SPECTSingle Photon Emission Computed Tomography
To increase efficiency:Most SPECT cameras have several detectors
From: IEEE TNS VOL. 42, NO. 4, 1995,Jingai Liu, Wei Chang, and Srecko Loncaric
Use fan-beam or cone-beam collimation when imaging small FOV (brain, pediatrics, cardiac..)
Dedicated systems for Brain and Cardiac imaging are considered.
Dedicated Brain SPECT systems
inSPira HD portable SPECT scanner for brain imaging from NeuroLogica corporation
inSPira HD features spiral-rotating focused collimators. Image quality approaches PET with the resulting reconstructed spatial resolution as high as 3.0mm. The focused collimators and spiral scan motion of the inSPira HD are responsible for the higher resolution (both in-plane and in z-axis) as compared to conventional Gamma Camera SPECT systems. The combination allows isotropic
scans and reconstruction with as high as 3mm resolution in X,Y,Z.
What is needed for SPECT?Complete set of projections for each axial plane.
1. The radiotracer distribution must is stable • The detector is always viewing the same
distribution• Patient is not move during the acquisition
2. Imaging system must be in alignment and have uniform and stable detectors
Assumptions:
Act
ivit
y in
Org
anTime
SPECT Imaging
YESNO
UniformRods SpheresComplete set of projections (360˚)
Incomplete set of projections (240˚)
Artifacts due to incomplete angular sampling:
Missing projections
180° acquisition is allowed only for Cardiac SPECT
due to anatomical position of the heart
Reconstructed region
Acquiring over 360 ° reduces some of the inconsistency associated with SPECT and reduces distortion, BUTTissue attenuation degrades the quality of the projections collected from the posteriorThe Nuclear Cardiology community has overwhelmingly endorsed 180 ° orbits
Effect of Collimator Damage (Nonuniformity) on Reconstructed Images
SPECT studybefore damage
SPECT studyafter damage
Integral Unif. = 2.5 %Differential Unif. = 1.9 %
Flood image of Head 2before damage
Integral Unif. = 4.2 %Differential Unif. = 3.4 %
Flood image of Head 2after damage
Damaged Collimator on Head 2
Quantitative SPECT?
The ultimate goal of quantitative SPECT is to provide reconstructedimages in which each pixel value in the image represents the absolute activity concentration in the corresponding region in the patient.
1. Collimator Blur2. Attenuation3. Scatter
The most important factors affecting quantitation are:
Source to Collimator distance vary as camera rotates during SPECT imaging.
FBP
Different Radial and Tangential Resolution
Tangential
Radial
GammaCamera
Collimator Blur:
Depth Dependant Resolution Recovery
From : M. O’Connor MWSNM-04 presentation
38 cm
28 cm
Average Size Patient
12.5 %
50.0 %
41 cm
37 cm
Large Patient
6.25 %
25.0%
Attenuation:
No Attenuation Correction:
Attenuation Correction Applied:
Accurate Attenuation Correction
•Iterative AlgorithmFilter backprojection cannot incorporate attenuation correction
•Co-registered Attenuation MapTransmission measurementsTransmission reconstructionAttenuation coefficient conversion
Multiple Point Sources (Philips)
Gantry mounted x-ray tube emits x-rays to an opposing CT detector (GE)
Emission + Transmission SPECTAttenuation Compensation
Multiple line sources in “wings” (Siemens)
Scanning line sources (Adac)
From : M. O’Connor MWSNM-04 presentation
From : M. O’Connor MWSNM-04 presentation
Attenuation Correction using Transmission Scan
UncorrectedCorrected
Scatter:• In a typical patient study with 99mTc labeled radiopharmaceutical, even using
narrow 15% PHA window, the ratio of the number of detected scattered photons to the number of nonscattered photons may be as large as 40%.
• The presence of scattered photons results in reduced image contrast and leads to an overestimation of the concentration of radioactivity in the pixel. The loss of image contrast may obscure clinically important details , particularly "cold" areas in the images.
Measured line profile
"Ideal" line profile
18014010060200Energy (keV)
Cou
nts
(arb
. u
nit
s)
15% window
Scattered photons
20 cm dia. cylinder filled with 99mTc and 6 cm dia. cold sphere
http:// http://www.physics.usyd.edu.au/ugrad/sphys/medphys.html
Scatter Correction Methods:
http:// http://www.physics.usyd.edu.au/ugrad/sphys/medphys.html
http://www.spectrum-dynamics.com/case_studies/comparison/patient_1.html
Detailed diagram of a single-detector column from D-SPECT camera.
Detector Specifications:Detector type: pixilated (16x64)Detection material: CdZnTeNumber of detectors: 9Field of view: 15.74cm x 3.94cmCollimator: built-in tungsten arrayEnergy resolution: <8% (99mTc)System uniformity: Integral: <4%Differential: <3%System Planar Sensitivity: 500 [cnts/μCi/min]Tungsten + Lead Shielding: 170keV
New dedicated SPECT systems:
D-SPECT
D-SPECT™
Conventional Camera:
16 min.
20 min.
D-SPECT:
2 min
4 min
Figure 6. Imaging with the CardiArc camera is accomplished via 3 curved NaI(Tl) crystals and an array of photomultiplier tubes. Collimation is achieved via a thin lead sheet with 6 vertical slits (aperture arc), which rotates during acquisition. (Courtesy of Dr. Jack Juni of CardiArc.)
From: Journal of Nuclear Cardiology, Patton, Slomka, and Berman, Volume 14, Number 4;501-13 Recent technologic advances in nuclear cardiology
Figure 7. Slice separation is accomplished with the CardiArc camera by use of thin lead vanes that are stacked vertically to define slices for imaging. (Courtesy of Dr. Jack Juni of CardiArc.)
CardiArc
Cardius 3 XPO triple-head, pixilated detector camera (Digirad).
Figure 3. Data acquisition with Cardius 3 XPO camera. The detectors remain fixed while the patient is rotated through 202.5° via a rotating chair configuration
Each detector head is 21.2 15.8 cm and contains an array of 768 6.1 6.1 5–mm thick CsI(Tl) crystals coupled to individual silicon photodiodes used to convert the light output of the crystals to electrical pulses. Digital Anger logic is used to process the signals and create images. In the 3-detector system the detector heads are fixed in position at 67.5° between heads. For imaging, the patient sits on a chair with his or her arms placed on an arm rest above the detectors. With this system, the manufacturer reports a reconstructed spatial resolution of 15.4 mm and a sensitivity of 234 cpm/Ci using the system’s cardiac collimator.
GE Healthcare - Alcyone Technology, a nuclear cardiology platform combining cadmium zinc telluride (CZT) detectors, focused pin-hole collimation, 3D reconstruction, and stationary data acquisition, to improve workflow, dose management, and overall image quality.
