Specs on kV imaging system
Varian G242 tube
14 degree anode
.4 and .8 mm focal spot
kvD amorphous silicon detector
Layer is 1.8 cm below the scatter grid
40x30 cm panel at isocenter
the kV arms are defined by anatomy, shoulder, arm etc
Defaults for the arms are 100 cm source to isocenter, 50 for the panel
kVp maxes out at 140, says it goes up to 150 but really 140.
OUT is out of the way, retract is fully retracted. But to access cabinets has to be OUT but not fully retracted (back).
Monday, December 29, 2008
Tuesday, September 16, 2008
Question and answer site
http://www.hps.org/publicinformation/ate/cat1.html
Lots of questions and answers from health physicists.
Lots of questions and answers from health physicists.
Monday, September 15, 2008
Bowtie filter
The field of view in the half-fan mode is 45 cm in
diameter with a 14 cm in the cranial caudal extent when SID
is 150 cm. An accessory, called a bowtie filter, is placed in
front of the kV beam to attenuate the edges of the kV beam.
The bowtie filter reduces skin dose, allows larger x-ray tech-
niques to be employed without saturating the detector, reduces
x-ray scatter, and reduces the effects of charge trapping
in the detector.
Source (Nov 2006 Medical Physics "A quality assurance program for the on board imager")
diameter with a 14 cm in the cranial caudal extent when SID
is 150 cm. An accessory, called a bowtie filter, is placed in
front of the kV beam to attenuate the edges of the kV beam.
The bowtie filter reduces skin dose, allows larger x-ray tech-
niques to be employed without saturating the detector, reduces
x-ray scatter, and reduces the effects of charge trapping
in the detector.
Source (Nov 2006 Medical Physics "A quality assurance program for the on board imager")
Tuesday, July 22, 2008
Thursday, July 17, 2008
Thursday, June 26, 2008
Dose Volume Constraints for Organs at Risk
Dose Volume Constraints for OARs
Conventional fractionation
Spinal Cord – 40 Gy (max)
Brainstem – 54 Gy (max)
Optic nerves – 54 Gy (max)
Parotid Glands – 35 Gy (max)
25 Gy (<50%)
Brain – 50 Gy (minimize volume above 30 Gy as much as possible)
Mandible – 70 Gy (max)
Cochlea/Middle Ear – 30 Gy (max)
Oral Cavity – 30 Gy (mean)
Brachial Plexus – 60 Gy (max)
Posterior Neck Avoidance – 45 Gy (mean)
Lens – 5 Gy (max)
Eyes – 45 Gy (max)
Optic Chiasm – 54 Gy (max)
Optice Nerves – 54 Gy (max)
Lung – 20 Gy (<35%)
Heart – 40 Gy (max)
Esophogus – 60 Gy (<50%)
Kidney – 20 Gy (max)
< 50% of combined both kidneys
< 75% of one side of kidney if another kidney is not functional
Liver – 30 Gy (<40%)
Femoral heads – 45 Gy (max)
Small Bowel – 48 Gy (max)
Rectum –
40 Gy (<60%)
45 Gy (<50%)
60 Gy (<40%)
70 Gy (<20%)
75.6 Gy (<15%)
78 Gy (<5%)
Bladder – 70 Gy (<20%)
Femoral Heads – 45 Gy (<50%)
Pituitary gland – 36 Gy (max, adults), 18 Gy (max, children), very safe
Dose Volume Constraints for OARs
SRS, single fraction
Brainstem – 12 Gy (max)
Optic nerves – 12 Gy (max)
Optic Chiasm – 12 Gy (max)
Optice Nerves – 12 Gy (max)
Retina – 12 Gy (max)
Normal Brain – 20 Gy (max)
Lens – 2 Gy (max)
Skin – 5 Gy (max)
Dose Volume Constraints for OARs
SBRT, 3-5 fractions
Spinal Cord – 18 Gy (max), 10 Gy (<10%) single fraction
3 Gy per fraction always safe, only draw cord within 4 mm of PTV
Lung – 20 Gy (<15% of total lung capacity) total
Heart – 6-8 Gy (max) per fraction
Esophogus – 6 Gy (max), 4 Gy (<75%) per fraction
Stomach – 8 Gy (max) per fraction
Small bowel – 4 Gy (max) per fraction
Kidney – 1 functioning kidney: 2 Gy (max), 1 Gy (<75%) per fraction
2 kidneys: 3 Gy (max) per fraction to one, test to make sure both work
Liver – 15 Gy (<700 cc) total
Bronchii and trachea – 8 Gy (max) per fraction
Skin – 4 Gy (max) per fraction
Conventional fractionation
Spinal Cord – 40 Gy (max)
Brainstem – 54 Gy (max)
Optic nerves – 54 Gy (max)
Parotid Glands – 35 Gy (max)
25 Gy (<50%)
Brain – 50 Gy (minimize volume above 30 Gy as much as possible)
Mandible – 70 Gy (max)
Cochlea/Middle Ear – 30 Gy (max)
Oral Cavity – 30 Gy (mean)
Brachial Plexus – 60 Gy (max)
Posterior Neck Avoidance – 45 Gy (mean)
Lens – 5 Gy (max)
Eyes – 45 Gy (max)
Optic Chiasm – 54 Gy (max)
Optice Nerves – 54 Gy (max)
Lung – 20 Gy (<35%)
Heart – 40 Gy (max)
Esophogus – 60 Gy (<50%)
Kidney – 20 Gy (max)
< 50% of combined both kidneys
< 75% of one side of kidney if another kidney is not functional
Liver – 30 Gy (<40%)
Femoral heads – 45 Gy (max)
Small Bowel – 48 Gy (max)
Rectum –
40 Gy (<60%)
45 Gy (<50%)
60 Gy (<40%)
70 Gy (<20%)
75.6 Gy (<15%)
78 Gy (<5%)
Bladder – 70 Gy (<20%)
Femoral Heads – 45 Gy (<50%)
Pituitary gland – 36 Gy (max, adults), 18 Gy (max, children), very safe
Dose Volume Constraints for OARs
SRS, single fraction
Brainstem – 12 Gy (max)
Optic nerves – 12 Gy (max)
Optic Chiasm – 12 Gy (max)
Optice Nerves – 12 Gy (max)
Retina – 12 Gy (max)
Normal Brain – 20 Gy (max)
Lens – 2 Gy (max)
Skin – 5 Gy (max)
Dose Volume Constraints for OARs
SBRT, 3-5 fractions
Spinal Cord – 18 Gy (max), 10 Gy (<10%) single fraction
3 Gy per fraction always safe, only draw cord within 4 mm of PTV
Lung – 20 Gy (<15% of total lung capacity) total
Heart – 6-8 Gy (max) per fraction
Esophogus – 6 Gy (max), 4 Gy (<75%) per fraction
Stomach – 8 Gy (max) per fraction
Small bowel – 4 Gy (max) per fraction
Kidney – 1 functioning kidney: 2 Gy (max), 1 Gy (<75%) per fraction
2 kidneys: 3 Gy (max) per fraction to one, test to make sure both work
Liver – 15 Gy (<700 cc) total
Bronchii and trachea – 8 Gy (max) per fraction
Skin – 4 Gy (max) per fraction
Friday, June 6, 2008
Mayneord F Factor
Mayneord Factor
* Overestimates the increase in PDD with increase in SSD
* Overestimates for small field sizes
* MF = ((f2+ dm) /(f1 + dm))squared x ((f1 + d) / (f2+ d))2
* Overestimates the increase in PDD with increase in SSD
* Overestimates for small field sizes
* MF = ((f2+ dm) /(f1 + dm))squared x ((f1 + d) / (f2+ d))2
PDD (Percentage Depth Dose) Key Points.
PDD
Increases with FS (less dependent with higher E)
Increases with E
Increases with SSD (due to ISL)
Decreases with Depth (exponentially) beyond build up region
Thursday, June 5, 2008
FILM: EDR VS XV
The two films are different in their response to dose:
XV films being to saturate at about 30 cGy
EDR2 films begin to saturate at about 300 cGy
The difference between the two films originates from their differences in the content of silver bromide crystals and grain size.
The grain size of EDR2 is smaller.
DOSIMETRIC PERFORMANCE
Compared with XV film, EDR2 film showed better agreement with calculations and measurements of dose. (olch2002)
EDR 2 film is less sensitive to low energy photons.
XV requires less amount of dose and thus the irradiation will be much quicker. On the other hand it will saturate much quicker than EDR2 film.
Source Film dosimetry (Yeo and Kim)
XV films being to saturate at about 30 cGy
EDR2 films begin to saturate at about 300 cGy
The difference between the two films originates from their differences in the content of silver bromide crystals and grain size.
The grain size of EDR2 is smaller.
DOSIMETRIC PERFORMANCE
Compared with XV film, EDR2 film showed better agreement with calculations and measurements of dose. (olch2002)
EDR 2 film is less sensitive to low energy photons.
XV requires less amount of dose and thus the irradiation will be much quicker. On the other hand it will saturate much quicker than EDR2 film.
Source Film dosimetry (Yeo and Kim)
Friday, May 30, 2008
Monday, May 19, 2008
Wednesday, May 14, 2008
Travelling vs Standing waves
· Accelerate electrons using a microwave using either a traveling wave, which absorbs the residual power with a dummy to prevent “backward reflected waves”, or a stationary wave accelerator which reflects the waves from each end towards each other to produce “standing waves”.
Mnemonic, when one is travelling they don't look backward (i.e. prevent backward reflecting waves)
When one is stationary or standing they can reflect back on their life (reflect waves to from each end)
Mnemonic, when one is travelling they don't look backward (i.e. prevent backward reflecting waves)
When one is stationary or standing they can reflect back on their life (reflect waves to from each end)
Sunday, May 11, 2008
Why use dref for electrons in TG-51
TG-51 gives a reference for the reference depth for electron beam, which is Ref.#17.Quote: For electron beam reference dosimetry in radiotherapy, it is shown that by choosingthe reference depth as dref=0.6R50 -0.1 cm, ..., the Spencer-Attix water-to-airstopping-power ratioat dref is given by (L/p)(water to air) = 1.2534-0.1487(R50)exp(0.2144).This is the magic.
REF YAHOO STUDY GROUP 4/23/08
REF YAHOO STUDY GROUP 4/23/08
Photon Interactions Part III
As far as the photon fate after the interaction with an
atom is concerned there are two possible outcomes:
• Photon disappears (i.e., is absorbed completely) and a portion
of its energy is transferred to light charged particles (electrons
and positrons in the absorbing medium).
• Photon is scattered and two outcomes are possible:
• The resulting photon has the same energy as the incident photon and no
light charged particles are released in the interaction.
• The resulting scattered photon has a lower energy than the incident photon
and the energy excess is transferred to a light charged particle (electron).
atom is concerned there are two possible outcomes:
• Photon disappears (i.e., is absorbed completely) and a portion
of its energy is transferred to light charged particles (electrons
and positrons in the absorbing medium).
• Photon is scattered and two outcomes are possible:
• The resulting photon has the same energy as the incident photon and no
light charged particles are released in the interaction.
• The resulting scattered photon has a lower energy than the incident photon
and the energy excess is transferred to a light charged particle (electron).
Photon Interaction (Tightly Bound)
1.4.1 Slide 5 (125/194)
1.4 PHOTON INTERACTIONS
1.4.1 Types of indirectly ionizing photon irradiations
A tightly bound electron is an electron whose binding
energy is comparable to, larger than, or slightly smaller
than the photon energy .
• For a photon interaction to occur with a tightly bound electron, the
binding energy of the electron must be of the order of, but
slightly smaller, than the photon energy
• An interaction between a photon and a tightly bound electron is
considered an interaction between photon and the atom as a
whole.
Photoelectric effect (tightly bound)
1.4 PHOTON INTERACTIONS
1.4.1 Types of indirectly ionizing photon irradiations
A tightly bound electron is an electron whose binding
energy is comparable to, larger than, or slightly smaller
than the photon energy .
• For a photon interaction to occur with a tightly bound electron, the
binding energy of the electron must be of the order of, but
slightly smaller, than the photon energy
• An interaction between a photon and a tightly bound electron is
considered an interaction between photon and the atom as a
whole.
Photoelectric effect (tightly bound)
Loosely bound electrons (Compton)
1.4 PHOTON INTERACTIONS
1.4.1 Types of indirectly ionizing photon irradiations
A loosely bound electron is an electron whose binding
energy to the nucleus is small compared to the
photon energy
An interaction between a photon and a loosely bound
electron is considered to be an interaction between a
photon and a free (unbound) electron.
1.4.1 Types of indirectly ionizing photon irradiations
A loosely bound electron is an electron whose binding
energy to the nucleus is small compared to the
photon energy
An interaction between a photon and a loosely bound
electron is considered to be an interaction between a
photon and a free (unbound) electron.
Photon Interactions part 2
1.4 PHOTON INTERACTIONS
1.4.1 Types of indirectly ionizing photon irradiations
Interactions of photons with nuclei may be:
• Direct photon-nucleus interactions (photodisintegration)
or
• Interactions between the photon and the electrostatic field of the
nucleus (pair production).
Photon-orbital electron interactions are characterized as
interactions between the photon and either
• A loosely bound electron (Compton effect, triplet production)
or
• A tightly bound electron (photoelectric effect).
1.4.1 Types of indirectly ionizing photon irradiations
Interactions of photons with nuclei may be:
• Direct photon-nucleus interactions (photodisintegration)
or
• Interactions between the photon and the electrostatic field of the
nucleus (pair production).
Photon-orbital electron interactions are characterized as
interactions between the photon and either
• A loosely bound electron (Compton effect, triplet production)
or
• A tightly bound electron (photoelectric effect).
Photon Interactions
1.4 PHOTON INTERACTIONS
1.4.1 Types of indirectly ionizing photon irradiations
In penetrating an absorbing medium, photons may
experience various interactions with the atoms of the
medium, involving:
• Absorbing atom as a whole
• Nuclei of the absorbing medium
• Orbital electrons of the absorbing medium.
1.4.1 Types of indirectly ionizing photon irradiations
In penetrating an absorbing medium, photons may
experience various interactions with the atoms of the
medium, involving:
• Absorbing atom as a whole
• Nuclei of the absorbing medium
• Orbital electrons of the absorbing medium.
Electron Interactions and Stopping Power
Electrons traversing an absorber lose their kinetic energy
through ionization collisions and radiation collisions.
The rate of energy loss per gram and per cm2 is called the
mass stopping power and it is a sum of two components:
• Mass collision stopping power
• Mass radiation stopping power
The rate of energy loss for a therapy electron beam in
water and water-like tissues, averaged over the electron’s
range, is about 2 MeV/cm.
through ionization collisions and radiation collisions.
The rate of energy loss per gram and per cm2 is called the
mass stopping power and it is a sum of two components:
• Mass collision stopping power
• Mass radiation stopping power
The rate of energy loss for a therapy electron beam in
water and water-like tissues, averaged over the electron’s
range, is about 2 MeV/cm.
Activity (Basics)
Activity represents the total number of disintegrations
(decays) of parent nuclei per unit time.
The SI unit of activity is the becquerel (1 Bq = 1 s-1).
Both the becquerel and the hertz correspond to s-1, however, hertz
expresses frequency of periodic motion, while becquerel expresses
activity.
The older unit of activity is the curie ,
originally defined as the activity of 1 g of radium-226.
Currently, the activity of 1 g of radium-226 is 0.988 Ci.
(1 Ci = 3.7 1010 s1)
(decays) of parent nuclei per unit time.
The SI unit of activity is the becquerel (1 Bq = 1 s-1).
Both the becquerel and the hertz correspond to s-1, however, hertz
expresses frequency of periodic motion, while becquerel expresses
activity.
The older unit of activity is the curie ,
originally defined as the activity of 1 g of radium-226.
Currently, the activity of 1 g of radium-226 is 0.988 Ci.
(1 Ci = 3.7 1010 s1)
Review of the basics Ch 1 Atomic Physics
The constituent particles forming an atom are:
• Proton
• Neutron
• Electron
Protons and neutrons are known as nucleons and they form the
nucleus.
Atomic number Z
Number of protons and number of electrons in an atom.
Atomic mass number A
Number of nucleons in an atom,
where
• Z is the number of protons (atomic number) in an atom.
• N is the number of neutrons in an atom.
• Proton
• Neutron
• Electron
Protons and neutrons are known as nucleons and they form the
nucleus.
Atomic number Z
Number of protons and number of electrons in an atom.
Atomic mass number A
Number of nucleons in an atom,
where
• Z is the number of protons (atomic number) in an atom.
• N is the number of neutrons in an atom.
Radiation fundamentals
Exposure (X), 1 R= 2.58 x10-4 C/kg air
Dose, 1Gy=100 rad
Equivalent Dose, 1 Sv=100 Rem (note weighting factors are applied)
Activity 1 Bq= 1Ci/3.7x10 to the power 10
Dose, 1Gy=100 rad
Equivalent Dose, 1 Sv=100 Rem (note weighting factors are applied)
Activity 1 Bq= 1Ci/3.7x10 to the power 10
Categories of ionizing radiation
Ionizing photon radiation is classified into four categories:
Characteristic x ray
Results from electronic transitions between atomic shells.
Bremsstrahlung
Results mainly from electron-nucleus Coulomb interactions.
Gamma ray
Results from nuclear transitions.
Annihilation quantum (annihilation radiation)
Results from positron-electron annihilation.
Courtesy Pgorsak Ch1
Characteristic x ray
Results from electronic transitions between atomic shells.
Bremsstrahlung
Results mainly from electron-nucleus Coulomb interactions.
Gamma ray
Results from nuclear transitions.
Annihilation quantum (annihilation radiation)
Results from positron-electron annihilation.
Courtesy Pgorsak Ch1
Saturday, May 10, 2008
Collisional Losses (Ionization and Excitation)
The rate of energy loss depends on the electron density of the medium.
The rate of energy loss per gram per centimeter squared, which is called the mass stopping power is greater for low atomic number (Z) material than for high Z materials (Khan pg 298). Compare the water to lead curve in Khan Fig 14.1
There are two reasons for this:
First, high Z materials have fewer electrons than low Z materials
Second, high Z have more tightly bound electrons, which are not as available for this type of interaction
The figure in Khan Ch 14 shows the the energy loss rate first decreases and then increases with increase in electron energy with a miniumum occurring at about 1 MeV. Above 1 MeV, the variation with energy is very gradual.
The energy loss rate of electrons of energy 1 MeV and above in water is roughly 2 MeV/cm.
The rate of energy loss per gram per centimeter squared, which is called the mass stopping power is greater for low atomic number (Z) material than for high Z materials (Khan pg 298). Compare the water to lead curve in Khan Fig 14.1
There are two reasons for this:
First, high Z materials have fewer electrons than low Z materials
Second, high Z have more tightly bound electrons, which are not as available for this type of interaction
The figure in Khan Ch 14 shows the the energy loss rate first decreases and then increases with increase in electron energy with a miniumum occurring at about 1 MeV. Above 1 MeV, the variation with energy is very gradual.
The energy loss rate of electrons of energy 1 MeV and above in water is roughly 2 MeV/cm.
Wednesday, May 7, 2008
Electron nucleus interactions
The energy loss by radiation and the radiative yield g increase directly
with the absorber atomic number Z and the kinetic energy of electrons.
The radiation yield for X ray targets in the diagnostic radiology energy
range (~100 keV) is of the order of 1%, while in the megavoltage energy
range it amounts to 10–20%.
with the absorber atomic number Z and the kinetic energy of electrons.
The radiation yield for X ray targets in the diagnostic radiology energy
range (~100 keV) is of the order of 1%, while in the megavoltage energy
range it amounts to 10–20%.
Indirectly Ionizing Radiation
Indirectly ionizing radiation (photons or neutrons) deposits energy in the
medium through a two step process:
● In the first step a charged particle is released in the medium (photons
release electrons or positrons, neutrons release protons or heavier ions);
● In the second step the released charged particles deposit energy to the
medium through direct Coulomb interactions with orbital electrons of the
atoms in the medium.
PGORSAK
medium through a two step process:
● In the first step a charged particle is released in the medium (photons
release electrons or positrons, neutrons release protons or heavier ions);
● In the second step the released charged particles deposit energy to the
medium through direct Coulomb interactions with orbital electrons of the
atoms in the medium.
PGORSAK
Saturday, May 3, 2008
Question 3542 Yahoo site, ADCL calibration
Can anyone tell me how ADCL calibrate user's ion chambers?ADCL website explains the configuration in the following.
Please see the diagram in the website.http://rpc.mdanderson.org/adcl/absorbed.htm"
Absorbed Dose to WaterAbsorbed dose to water calibrations are performed in a 30 x 30 x 30cm3 phantom at 5 cm depth and 85 cm from the source. Water proofchambers are calibrated bare in water, other chambers are calibrated in1 mm PMMA water-proofing provided by the ADCL. The field size atthe location of the chamber is 10 x 10 cm2. The dose rate at thelocation of the chamber will be between 25 and 50 cGy/min."
I have a question here.TG-51 says that "The gradienteffects are included implicitly in the beam quality conversionfactor kQ for photons and explicitly by the term PgrQ for electrons."How the gradient effect is included for x-ray?I think it is something to do with how ADCL calibrate user's ionchamber.Can anyone explain?
Thanks,Jongmin
Please see the diagram in the website.http://rpc.mdanderson.org/adcl/absorbed.htm"
Absorbed Dose to WaterAbsorbed dose to water calibrations are performed in a 30 x 30 x 30cm3 phantom at 5 cm depth and 85 cm from the source. Water proofchambers are calibrated bare in water, other chambers are calibrated in1 mm PMMA water-proofing provided by the ADCL. The field size atthe location of the chamber is 10 x 10 cm2. The dose rate at thelocation of the chamber will be between 25 and 50 cGy/min."
I have a question here.TG-51 says that "The gradienteffects are included implicitly in the beam quality conversionfactor kQ for photons and explicitly by the term PgrQ for electrons."How the gradient effect is included for x-ray?I think it is something to do with how ADCL calibrate user's ionchamber.Can anyone explain?
Thanks,Jongmin
Thursday, May 1, 2008
Neutron shielding
Ch 5 Mcginley pg 69
Most medical accelerators operating above 10 MeV use a maze with a door shielded for neutrons and photons at the outer maze entrance. A typical door consists of a steel case 0.635 cm thick containing 10.2 cm of borated polyethylene.
The polyethylene is used to moderate the fast and intermediate energy neutrons, which then react with the boron and produce a 0.473 MeV photon.
The lead is placed after the polyethylene where it will attenuate the photons where it will attenuate the photons produced in the boron and any capture gamma rays generated in the maze by neutron capture in the concrete in the concrete wall, ceiling and floor.
Recently McCall (1997) has indicated a more efficient door shield is produced by placing the lead before the polyethylene. With this arrangement, the high energy neutron component is reduced in energy by interactions in the lead before entering the polyethylene layer.
The overall result is an increased attenuation of the neutrons with approximately the same penetration for the capture gamma rays that originate in the maze. The low energy photons aristing from the (n, alpha) reaction with the boron in the polyethylene are attenuated sufficiently by the steel case of the door.
Most medical accelerators operating above 10 MeV use a maze with a door shielded for neutrons and photons at the outer maze entrance. A typical door consists of a steel case 0.635 cm thick containing 10.2 cm of borated polyethylene.
The polyethylene is used to moderate the fast and intermediate energy neutrons, which then react with the boron and produce a 0.473 MeV photon.
The lead is placed after the polyethylene where it will attenuate the photons where it will attenuate the photons produced in the boron and any capture gamma rays generated in the maze by neutron capture in the concrete in the concrete wall, ceiling and floor.
Recently McCall (1997) has indicated a more efficient door shield is produced by placing the lead before the polyethylene. With this arrangement, the high energy neutron component is reduced in energy by interactions in the lead before entering the polyethylene layer.
The overall result is an increased attenuation of the neutrons with approximately the same penetration for the capture gamma rays that originate in the maze. The low energy photons aristing from the (n, alpha) reaction with the boron in the polyethylene are attenuated sufficiently by the steel case of the door.
Gadolinium for MRI
Breast MRI Contrast Agent: Gadolinium
Malignant breast tumors begin to grow their own blood supply network once they reach a certain size; this is the only way the cancer can continue to grow. In a breast MRI scan, a contrast agent injected into the bloodstream can provide information about blood supply to the breast tissues; the agent "lights up" a tumor by highlighting its blood vessel network. Usually, several scans are taken: one before the contrast agent is injected and at least one after. The pre-contrast and post-contrast images are compared and areas of difference are highlighted. It is important to note that if the patient moves even slightly between the two scans, the shape or size of the image may be distorted--a big loss of information.
What is Gadolinium?
This is an FDA approved contrast agent for MRI. Gadolinium, or gadodiamide, provides greater contrast between normal tissue and abnormal tissue in the brain and body. Gadolinium looks clear like water and is non-radioactive. After it is injected into a vein, Gadolinium accumulates in the abnormal tissue that may be affecting the body or head. Gadolinium causes these abnormal areas to become very bright (enhanced) on the MRI. This makes it very easy to see. Gadolinium is then rapidly cleared from the body by the kidneys.
Does Gadolinium go by other names?
Gadolinium, gadolinium-DPTA, gadodiamide. It also goes by various brand names, depending on the pharmaceutical company that makes it:
Magnevist (Berlex Laboratories, Inc.)
Omniscan (Nycomed Amersham plc)
ProHance (Bracco Diagnostics, Inc.)
What does Gadolinium do?
Gadolinium allows the MRI to define abnormal tissue with greater clarity than ever before. Tumors enhance after Gadolinium is given. The exact size of the tumor and location are very important in treatment planning and follow up. Gadolinium is also helpful in finding small tumors by making them bright and easy to see.
Is Gadolinium safe?
Gadolinium has been used for years in adults and children in the United States, Europe and Japan, without any serious complications in thousands of patients. The FDA declared Gadolinium safe for use in MRI in 1988. A few side effects, such as mild headache, nausea and local burning can occur. Very rarely (less than one in a thousand), patients are allergic to Gadolinium. If you have kidney problems, it must be used with caution. Gadolinium should be used in pregnant patients or nursing mothers only when the benefits outweigh the risk. Gadolinium used in MRI is many times safer than the iodine type contrast used in CT scans. There is more information at the International MR Safety Central Web Site.
What are the side effects?
Side effects of the contrast agent injection include mild headache, nausea and local pain. Rarely (less than 1% of the time) low blood pressure and lightheadedness occurs. This can be treated immediately with intravenous fluids. Very rarely (less than one in one thousand), patients are allergic to the contrast agent. These effects are most commonly hives and itchy eyes, but more severe reactions have been seen which result in shortness of breath.
What about breast feeding?
According to Dr. Emanual Kanal of the International MR Safety Central Web Site: "More data are available for Magnevist than for the other agents. Magnevist is excreted in very low concentrations (i.e., 0.011% of the total dose) in human breast milk over approximately 33 hours. The concentration of this contrast agent in breast milk peaks at approximately 4.75 hours and decreases to less than a fifth of this level (to less than 1 micromol/L) 22 hours after injection. For this reason, and as an extra precaution, it is recommended that nursing mothers express their breasts and not breastfeed for 36 to 48 hours after administration of an MR imaging contrast agent, to ensure that the nursing child does not receive the drug in any appreciable quantity."
What about new contrast agents?
New contrast agents to improve breast MRI are currently being developed and studied in safety and efficacy clinical trials.
Malignant breast tumors begin to grow their own blood supply network once they reach a certain size; this is the only way the cancer can continue to grow. In a breast MRI scan, a contrast agent injected into the bloodstream can provide information about blood supply to the breast tissues; the agent "lights up" a tumor by highlighting its blood vessel network. Usually, several scans are taken: one before the contrast agent is injected and at least one after. The pre-contrast and post-contrast images are compared and areas of difference are highlighted. It is important to note that if the patient moves even slightly between the two scans, the shape or size of the image may be distorted--a big loss of information.
What is Gadolinium?
This is an FDA approved contrast agent for MRI. Gadolinium, or gadodiamide, provides greater contrast between normal tissue and abnormal tissue in the brain and body. Gadolinium looks clear like water and is non-radioactive. After it is injected into a vein, Gadolinium accumulates in the abnormal tissue that may be affecting the body or head. Gadolinium causes these abnormal areas to become very bright (enhanced) on the MRI. This makes it very easy to see. Gadolinium is then rapidly cleared from the body by the kidneys.
Does Gadolinium go by other names?
Gadolinium, gadolinium-DPTA, gadodiamide. It also goes by various brand names, depending on the pharmaceutical company that makes it:
Magnevist (Berlex Laboratories, Inc.)
Omniscan (Nycomed Amersham plc)
ProHance (Bracco Diagnostics, Inc.)
What does Gadolinium do?
Gadolinium allows the MRI to define abnormal tissue with greater clarity than ever before. Tumors enhance after Gadolinium is given. The exact size of the tumor and location are very important in treatment planning and follow up. Gadolinium is also helpful in finding small tumors by making them bright and easy to see.
Is Gadolinium safe?
Gadolinium has been used for years in adults and children in the United States, Europe and Japan, without any serious complications in thousands of patients. The FDA declared Gadolinium safe for use in MRI in 1988. A few side effects, such as mild headache, nausea and local burning can occur. Very rarely (less than one in a thousand), patients are allergic to Gadolinium. If you have kidney problems, it must be used with caution. Gadolinium should be used in pregnant patients or nursing mothers only when the benefits outweigh the risk. Gadolinium used in MRI is many times safer than the iodine type contrast used in CT scans. There is more information at the International MR Safety Central Web Site.
What are the side effects?
Side effects of the contrast agent injection include mild headache, nausea and local pain. Rarely (less than 1% of the time) low blood pressure and lightheadedness occurs. This can be treated immediately with intravenous fluids. Very rarely (less than one in one thousand), patients are allergic to the contrast agent. These effects are most commonly hives and itchy eyes, but more severe reactions have been seen which result in shortness of breath.
What about breast feeding?
According to Dr. Emanual Kanal of the International MR Safety Central Web Site: "More data are available for Magnevist than for the other agents. Magnevist is excreted in very low concentrations (i.e., 0.011% of the total dose) in human breast milk over approximately 33 hours. The concentration of this contrast agent in breast milk peaks at approximately 4.75 hours and decreases to less than a fifth of this level (to less than 1 micromol/L) 22 hours after injection. For this reason, and as an extra precaution, it is recommended that nursing mothers express their breasts and not breastfeed for 36 to 48 hours after administration of an MR imaging contrast agent, to ensure that the nursing child does not receive the drug in any appreciable quantity."
What about new contrast agents?
New contrast agents to improve breast MRI are currently being developed and studied in safety and efficacy clinical trials.
Wednesday, April 30, 2008
Monday, April 28, 2008
Collapsed Cone Convolution Superposition (from ADAC)
The ADAC Pinnacle3 Collapsed Cone Convolution Superposition Dose Model
Todd McNutt, Ph.D. – Director of Product Development
ADAC’s Pinnacle3 3D treatment planning system uses a Collapsed Cone Convolution Superposition computation to determine the dose distribution in patients from external photon beams. The Pinnacle3 Convolution Superposition dose model is a true three-dimensional dose computation which intrinsically handles the effects of patient heterogeneities on both primary and secondary scattered radiation. This computation method is uniquely able to account for dose distributions in areas where the electronic equilibrium is perturbed, such as tissue-air interfaces and tissue-bone interfaces. While other convolution techniques account for the effects of patient heterogeneities on primary radiation, they neglect the effects of heterogeneities on scattered radiation in the final dose distribution. In addition, the nature of the Convolution Superposition dose model makes it ideally suited to handle optimization and intensity modulated radiation therapy planning.
The Convolution Superposition Dose Model
The Pinnacle3 Convolution Superposition dose algorithm is based on the work of Mackie, et al. Rather than correcting measured dose distributions, the Convolution Superposition algorithm computes dose distributions from first principles and, therefore, can account for the effects of beam modifiers, the external patient contour, and tissue heterogeneities on the dose distribution.
The Convolution Superposition dose model consists of four parts:
• Modeling the incident energy fluence as it exits the accelerator head. • Projection of this energy fluence through the density representation of a patient to compute a TERMA (Total Energy Released per unit Mass) volume. • A three-dimensional superposition of the TERMA with an energy deposition kernel using a ray-tracing technique to incorporate the effects of heterogeneities on lateral scatter. • Electron contamination is modeled with an exponential falloff which is added to the dose distribution after the photon dose is computed. The following sections describe each part of the model in more detail.
Modeling the Incident Energy Fluence as it exits the Accelerator
The incident energy fluence distribution is modeled as a two-dimensional array which describes the radiation exiting the head of the linear accelerator. The parameters defining this array are defined during physics data modeling.
The starting point for photon modeling is a uniform plane of energy fluence describing the intensity of the radiation exiting the accelerator head. The fluence model is then adjusted to account for the flattening filter, the accelerator head, and beam modifiers such as blocks, wedges and compensators.
• The “horns” in the beam produced by the flattening filter are modeled by removing an inverted cone from the distribution. • Off-focus scatter produced in the accelerator head is modeled by defining a 2D Gaussian function as a scatter source and adjusting the incident energy fluence based on the portion of the Gaussian distribution visible from each point in the incident energy fluence plane. • The geometric penumbra is modeled by convolving the array with a focal spot blurring function. • During planning, the shape of the field produced by blocks or multi-leaf collimators is cut out of the array leaving behind the corresponding transmission through the shape-defining entity. • Beam modifiers such as wedges and compensators are included in the array by attenuating the energy fluence by the corresponding thickness of the modifier. For static wedges and compensators, a radiological depth array is also stored which allows for proper modeling of the beam hardening due to the presence of the beam modifiers during the projection of the incident fluence array. Dynamic beam delivery with intensity-modulation or dynamic wedges is easily handled using the incident energy fluence array. For these beams, the radiological depth array is not needed to account for beam hardening.
Projection of Energy Fluence through a CT Patient Representation
The incident energy fluence plane is projected through the CT patient representation and attenuated using mass attenuation coefficients. These coefficients are stored in a three-dimensional lookup table as a function of density, radiological depth, and off-axis angle. Patient heterogeneities are taken into account with the density dependence. Beam hardening through the patient is accounted for with the radiological depth dependence, and the off-axis softening of the energy spectrum is produced with the off-axis angle dependence. To account for the changes in the photon energy spectrum at different locations in the beam, the mass attenuation coefficient lookup table is produced using a weighted sum of several mono-energetic tables.
The TERMA (Total Energy Released per unit Mass) volume is computed by projecting the incident energy fluence through the patient density volume using a ray-tracing technique. A given ray’s direction is determined based on the position of the radiation source and the particular location in the incident fluence plane. At each voxel in the ray path, the TERMA is computed using the attenuated energy fluence along the ray and the mass attenuation coefficient at the particular density, radiological depth, and off-axis angle.
3D Superposition of an Energy Deposition Kernel
The three-dimensional dose distribution in the patient is computed by superposition of the TERMA volume with the energy deposition kernel. The kernel represents the spread of energy from the primary photon interaction site throughout the associated volume. Poly-energetic kernels are produced by combining a series of Monte Carlo-generated mono-energetic energy deposition kernels. The superposition is carried out using a ray tracing technique similar to that used in the projection of the incident energy fluence. The kernel is inverted so that the dose can be computed in only a portion of the patient (TERMA) volume if desired. This allows for point dose computation and decreases computation time.
The rays from the dose deposition site are cast in three dimensions. At each voxel of the TERMA traversed along a ray, the contribution of dose to the dose deposition site is computed and accumulated using the TERMA and the kernel value at the current radiological distance. Using the radiological distance along the ray also allows the kernel to be scaled to account for the presence of heterogeneities with respect to scattered radiation in all directions.
The dose computation described above determines the dose from a single beam. Multiple beams are computed independently and the entire 3D dose distribution is created by adding the dose from each beam together according to the corresponding beam weight.
Adaptive Convolution Superposition
An Adaptive Convolution Superposition approach has also been implemented in Pinnacle3. This uses the calculation technique described above with some slight modifications. The speed of the computation is increased by adaptively varying the resolution of the dose computation grid depending on the curvature of the TERMA and dose distribution. First, the dose in a coarse 3D grid is computed and then the curvature in the TERMA distribution is assessed. In regions where the curvature is high, the dose is computed at intermediate points to provide higher resolution. The system adaptively increases the resolution in regions of high curvature until an acceptable resolution is used. In regions of low curvature, the dose is interpolated from the coarse dose grid. This technique decreases the computation time by a factor of 2-3 without compromising the accuracy of the Convolution Superposition calculation in the presence of heterogeneities.
Other Model-Based Algorithms
Other model-based algorithms, including 3D Fast Fourier Transform (FFT) techniques and differential pencil beam models which use FFTs on two-dimensional planes, use a projection of the incident energy fluence similar to that used in Pinnacle3 to determine the TERMA volume. They differ in that they do not use the superposition technique in the convolution process.
The FFT techniques require the assumption of an invariant kernel, which inherently assumes a homogeneous density representation during the convolution process. This technique reduces the accuracy of the computation because it ignores the effects of heterogeneities on laterally scattered radiation. Some post-computation corrections may be performed to help alleviate the error. In contrast, by scaling the rays from the primary dose deposition site, the Convolution Superposition method accurately and intrinsically models the effects of lateral scatter from tissue heterogeneities, a requirement for calculating dose from conformal and intensity modulated fields.
Although the FFT algorithms have a fast computation speed per computation point, they require full computation of dose over the entire TERMA volume. For irregular field calculations, point doses, plane doses, or other situations where the planner is only interested in the dose to smaller regions, the FFT algorithms still require the dose to be computed over the entire volume. The Convolution Superposition method can accurately compute the dose to a single point with the inverted energy deposition kernel. Therefore, depending on the planning situation, the desired dose calculation may be faster using the Convolution Superposition calculation than when using the FFT calculation.
The ability to define a smaller calculation matrix also results in a calculation speed advantage with the Convolution Superposition model for inverse planning and optimization of intensity modulated beams where computation of dose need only be performed in limited regions during the optimization process.
Further Reading
This paper provides an overview of the Convolution Superposition dose computation method used in the ADAC Pinnacle3 3D treatment planning system. For further information on this method and other dose computation techniques, please refer to the publications listed below:
R. Mohan, C. Chui, L. Lidofsky, “Energy and angular distributions of photons from medical linear accelerators,” Med. Phys. 12, 592-597 (1985). T.R. Mackie, J.W. Scrimger, J.J. Battista, “A convolution method of calculating dose for 15-MV x-rays,” Med. Phys. 12, 188-196 (1985). T.R. Mackie, A. Ahnesjo, P. Dickof, A. Snider, “Development of a convolution/superposition method for photon beams,” Use of Comp. In Rad. Ther., 107-110 (1987). A. Ahnesjo, P. Andreo, A. Brahme, “Calculation and application of point spread functions for treatment planning with high energy photon beams,” Acta. Oncol., 26, 49-56, (1987). T.R. Mackie, A.F. Bielajew, D.W.O. Rogers, J.J. Battista, “Generation of photon energy deposition kernels using the EGS Monte Carlo code,” Phys. Med. Biol. 33, 1-20 (1988). T.R. Mackie, P.J. Reckwerdt, T.W. Holmes, S.S. Kubsad, “Review of convolution/superposition methods for photon beam dose computation,” Proceedings of the Xth ICCR, 20-23, (1990). T.R. Mackie, P.J. Reckwerdt, M.A. Gehring, T.W. Holmes, S.S. Kubsad, B.R. Thomadsen, C.A. Sanders, B.R. Paliwal, T.J. Kinsella, “Clinical implementation of the convolution/superposition method,” Proceedings of the Xth ICCR, 322-325, (1990). N. Papanikolaou, T.R. Mackie, C. Meger-Wells, M. Gehring, P. Reckwerdt, “Investigation of the convolution method for polyenergetic spectra,” Med. Phys. 20, 1327-1336 (1993). M.B. Sharpe, J.J. Battista, “Dose calculations using convolution and superposition principles: The orientation of the dose spread kernels in divergent x-ray beams,” Med. Phys., 20, 1685-1694 (1993). T.R. McNutt, T.R. Mackie, P. Reckwerdt, N. Papanikolaou, B.R. Paliwal, “Calculation of portal dose images using the convolution/superposition method,” Med. Phys. 23(4) (1996). T.R. Mackie, P.J. Reckwerdt, T.R. McNutt, M. Gehring, C. Sanders, “Photon dose computations,” Teletherapy: Proceedings of the 1996 AAPM Summer School, Ed. J. Palta, T. R. Mackie., AAPM-College Park, MD, (1996). ADAC Laboratories 540 Alder Dr. Milpitas, CA 95035 Tel: (408)321-9100 (800)538-8531 Fax: (408)577-0907 www.adaclabs.com
MBA-X0010, Rev. B
Todd McNutt, Ph.D. – Director of Product Development
ADAC’s Pinnacle3 3D treatment planning system uses a Collapsed Cone Convolution Superposition computation to determine the dose distribution in patients from external photon beams. The Pinnacle3 Convolution Superposition dose model is a true three-dimensional dose computation which intrinsically handles the effects of patient heterogeneities on both primary and secondary scattered radiation. This computation method is uniquely able to account for dose distributions in areas where the electronic equilibrium is perturbed, such as tissue-air interfaces and tissue-bone interfaces. While other convolution techniques account for the effects of patient heterogeneities on primary radiation, they neglect the effects of heterogeneities on scattered radiation in the final dose distribution. In addition, the nature of the Convolution Superposition dose model makes it ideally suited to handle optimization and intensity modulated radiation therapy planning.
The Convolution Superposition Dose Model
The Pinnacle3 Convolution Superposition dose algorithm is based on the work of Mackie, et al. Rather than correcting measured dose distributions, the Convolution Superposition algorithm computes dose distributions from first principles and, therefore, can account for the effects of beam modifiers, the external patient contour, and tissue heterogeneities on the dose distribution.
The Convolution Superposition dose model consists of four parts:
• Modeling the incident energy fluence as it exits the accelerator head. • Projection of this energy fluence through the density representation of a patient to compute a TERMA (Total Energy Released per unit Mass) volume. • A three-dimensional superposition of the TERMA with an energy deposition kernel using a ray-tracing technique to incorporate the effects of heterogeneities on lateral scatter. • Electron contamination is modeled with an exponential falloff which is added to the dose distribution after the photon dose is computed. The following sections describe each part of the model in more detail.
Modeling the Incident Energy Fluence as it exits the Accelerator
The incident energy fluence distribution is modeled as a two-dimensional array which describes the radiation exiting the head of the linear accelerator. The parameters defining this array are defined during physics data modeling.
The starting point for photon modeling is a uniform plane of energy fluence describing the intensity of the radiation exiting the accelerator head. The fluence model is then adjusted to account for the flattening filter, the accelerator head, and beam modifiers such as blocks, wedges and compensators.
• The “horns” in the beam produced by the flattening filter are modeled by removing an inverted cone from the distribution. • Off-focus scatter produced in the accelerator head is modeled by defining a 2D Gaussian function as a scatter source and adjusting the incident energy fluence based on the portion of the Gaussian distribution visible from each point in the incident energy fluence plane. • The geometric penumbra is modeled by convolving the array with a focal spot blurring function. • During planning, the shape of the field produced by blocks or multi-leaf collimators is cut out of the array leaving behind the corresponding transmission through the shape-defining entity. • Beam modifiers such as wedges and compensators are included in the array by attenuating the energy fluence by the corresponding thickness of the modifier. For static wedges and compensators, a radiological depth array is also stored which allows for proper modeling of the beam hardening due to the presence of the beam modifiers during the projection of the incident fluence array. Dynamic beam delivery with intensity-modulation or dynamic wedges is easily handled using the incident energy fluence array. For these beams, the radiological depth array is not needed to account for beam hardening.
Projection of Energy Fluence through a CT Patient Representation
The incident energy fluence plane is projected through the CT patient representation and attenuated using mass attenuation coefficients. These coefficients are stored in a three-dimensional lookup table as a function of density, radiological depth, and off-axis angle. Patient heterogeneities are taken into account with the density dependence. Beam hardening through the patient is accounted for with the radiological depth dependence, and the off-axis softening of the energy spectrum is produced with the off-axis angle dependence. To account for the changes in the photon energy spectrum at different locations in the beam, the mass attenuation coefficient lookup table is produced using a weighted sum of several mono-energetic tables.
The TERMA (Total Energy Released per unit Mass) volume is computed by projecting the incident energy fluence through the patient density volume using a ray-tracing technique. A given ray’s direction is determined based on the position of the radiation source and the particular location in the incident fluence plane. At each voxel in the ray path, the TERMA is computed using the attenuated energy fluence along the ray and the mass attenuation coefficient at the particular density, radiological depth, and off-axis angle.
3D Superposition of an Energy Deposition Kernel
The three-dimensional dose distribution in the patient is computed by superposition of the TERMA volume with the energy deposition kernel. The kernel represents the spread of energy from the primary photon interaction site throughout the associated volume. Poly-energetic kernels are produced by combining a series of Monte Carlo-generated mono-energetic energy deposition kernels. The superposition is carried out using a ray tracing technique similar to that used in the projection of the incident energy fluence. The kernel is inverted so that the dose can be computed in only a portion of the patient (TERMA) volume if desired. This allows for point dose computation and decreases computation time.
The rays from the dose deposition site are cast in three dimensions. At each voxel of the TERMA traversed along a ray, the contribution of dose to the dose deposition site is computed and accumulated using the TERMA and the kernel value at the current radiological distance. Using the radiological distance along the ray also allows the kernel to be scaled to account for the presence of heterogeneities with respect to scattered radiation in all directions.
The dose computation described above determines the dose from a single beam. Multiple beams are computed independently and the entire 3D dose distribution is created by adding the dose from each beam together according to the corresponding beam weight.
Adaptive Convolution Superposition
An Adaptive Convolution Superposition approach has also been implemented in Pinnacle3. This uses the calculation technique described above with some slight modifications. The speed of the computation is increased by adaptively varying the resolution of the dose computation grid depending on the curvature of the TERMA and dose distribution. First, the dose in a coarse 3D grid is computed and then the curvature in the TERMA distribution is assessed. In regions where the curvature is high, the dose is computed at intermediate points to provide higher resolution. The system adaptively increases the resolution in regions of high curvature until an acceptable resolution is used. In regions of low curvature, the dose is interpolated from the coarse dose grid. This technique decreases the computation time by a factor of 2-3 without compromising the accuracy of the Convolution Superposition calculation in the presence of heterogeneities.
Other Model-Based Algorithms
Other model-based algorithms, including 3D Fast Fourier Transform (FFT) techniques and differential pencil beam models which use FFTs on two-dimensional planes, use a projection of the incident energy fluence similar to that used in Pinnacle3 to determine the TERMA volume. They differ in that they do not use the superposition technique in the convolution process.
The FFT techniques require the assumption of an invariant kernel, which inherently assumes a homogeneous density representation during the convolution process. This technique reduces the accuracy of the computation because it ignores the effects of heterogeneities on laterally scattered radiation. Some post-computation corrections may be performed to help alleviate the error. In contrast, by scaling the rays from the primary dose deposition site, the Convolution Superposition method accurately and intrinsically models the effects of lateral scatter from tissue heterogeneities, a requirement for calculating dose from conformal and intensity modulated fields.
Although the FFT algorithms have a fast computation speed per computation point, they require full computation of dose over the entire TERMA volume. For irregular field calculations, point doses, plane doses, or other situations where the planner is only interested in the dose to smaller regions, the FFT algorithms still require the dose to be computed over the entire volume. The Convolution Superposition method can accurately compute the dose to a single point with the inverted energy deposition kernel. Therefore, depending on the planning situation, the desired dose calculation may be faster using the Convolution Superposition calculation than when using the FFT calculation.
The ability to define a smaller calculation matrix also results in a calculation speed advantage with the Convolution Superposition model for inverse planning and optimization of intensity modulated beams where computation of dose need only be performed in limited regions during the optimization process.
Further Reading
This paper provides an overview of the Convolution Superposition dose computation method used in the ADAC Pinnacle3 3D treatment planning system. For further information on this method and other dose computation techniques, please refer to the publications listed below:
R. Mohan, C. Chui, L. Lidofsky, “Energy and angular distributions of photons from medical linear accelerators,” Med. Phys. 12, 592-597 (1985). T.R. Mackie, J.W. Scrimger, J.J. Battista, “A convolution method of calculating dose for 15-MV x-rays,” Med. Phys. 12, 188-196 (1985). T.R. Mackie, A. Ahnesjo, P. Dickof, A. Snider, “Development of a convolution/superposition method for photon beams,” Use of Comp. In Rad. Ther., 107-110 (1987). A. Ahnesjo, P. Andreo, A. Brahme, “Calculation and application of point spread functions for treatment planning with high energy photon beams,” Acta. Oncol., 26, 49-56, (1987). T.R. Mackie, A.F. Bielajew, D.W.O. Rogers, J.J. Battista, “Generation of photon energy deposition kernels using the EGS Monte Carlo code,” Phys. Med. Biol. 33, 1-20 (1988). T.R. Mackie, P.J. Reckwerdt, T.W. Holmes, S.S. Kubsad, “Review of convolution/superposition methods for photon beam dose computation,” Proceedings of the Xth ICCR, 20-23, (1990). T.R. Mackie, P.J. Reckwerdt, M.A. Gehring, T.W. Holmes, S.S. Kubsad, B.R. Thomadsen, C.A. Sanders, B.R. Paliwal, T.J. Kinsella, “Clinical implementation of the convolution/superposition method,” Proceedings of the Xth ICCR, 322-325, (1990). N. Papanikolaou, T.R. Mackie, C. Meger-Wells, M. Gehring, P. Reckwerdt, “Investigation of the convolution method for polyenergetic spectra,” Med. Phys. 20, 1327-1336 (1993). M.B. Sharpe, J.J. Battista, “Dose calculations using convolution and superposition principles: The orientation of the dose spread kernels in divergent x-ray beams,” Med. Phys., 20, 1685-1694 (1993). T.R. McNutt, T.R. Mackie, P. Reckwerdt, N. Papanikolaou, B.R. Paliwal, “Calculation of portal dose images using the convolution/superposition method,” Med. Phys. 23(4) (1996). T.R. Mackie, P.J. Reckwerdt, T.R. McNutt, M. Gehring, C. Sanders, “Photon dose computations,” Teletherapy: Proceedings of the 1996 AAPM Summer School, Ed. J. Palta, T. R. Mackie., AAPM-College Park, MD, (1996). ADAC Laboratories 540 Alder Dr. Milpitas, CA 95035 Tel: (408)321-9100 (800)538-8531 Fax: (408)577-0907 www.adaclabs.com
MBA-X0010, Rev. B
Sunday, April 27, 2008
Bone and Co-60
Presence of bone in the beam path decreases the dose behind bone for a Co-60 beam by about what percentage?
3.5%
3.5%
Bone and 10 MV photons
Presence of bone in the beam path decreases the dose behind bone (per cm of bone) for a 10 MV x-ray beam by about 2%
Bone and heterogeneity corrections
Two important effects due to the presence of bone in the patient that must be taken into account during treatment planning with electron beams are
1) reduction in the dose behind the bone
2) lateral scattering of electrons in bone giving rise to a high dose region in the patient.
1) reduction in the dose behind the bone
2) lateral scattering of electrons in bone giving rise to a high dose region in the patient.
Lung and inhomogeneity corrections
The presence of lung in the beam path increases the dose behind the lung (per cm of lung tissue), for an 10 MV x-ray beam, by approximately
2% per cm increase in dose behind lung
The presence of lung in the beam path increases the dose behind the Co-60 beam by approximately 4%
2% per cm increase in dose behind lung
The presence of lung in the beam path increases the dose behind the Co-60 beam by approximately 4%
Saturday, April 26, 2008
Tuesday, April 15, 2008
TG-40 Annual
TG-40
Annual
Dosimetry Tolerance
X-Ray Output Calibration Constancy 2%
Field size dependence of x-ray output constancy 2%
Output factor constancy for electron applicators 2%
Central axis parameter constancy (PDD.TAR) 2%
Off axis factor constancy 2%
Transmission factor constancy for all treatment accessories 2%
Wedge transmission factor constancy** 2%
Monitor chamber linearity 2%
X-ray output constancy vs gantry angle 2%
Electron Output Constancy vs gantry angle 2%
Off axis factor constancy vs gantry angle 2%
Arc mode Mfrs specs
Mechanical Checks
Collimator rotation isocenter 2 mm diameter
Gantry rotation isocenter 2 mm diameter
Couch rotation isocenter 2 mm diameter
Coincidence of collimetry,gantry, couch axes with isocenter 2 mm diameter
Coincidence of radiation and mechanical isocenter 2mm diameter
Table top sag 2mm
Vertical travel of table 2mm
Annual
Dosimetry Tolerance
X-Ray Output Calibration Constancy 2%
Field size dependence of x-ray output constancy 2%
Output factor constancy for electron applicators 2%
Central axis parameter constancy (PDD.TAR) 2%
Off axis factor constancy 2%
Transmission factor constancy for all treatment accessories 2%
Wedge transmission factor constancy** 2%
Monitor chamber linearity 2%
X-ray output constancy vs gantry angle 2%
Electron Output Constancy vs gantry angle 2%
Off axis factor constancy vs gantry angle 2%
Arc mode Mfrs specs
Mechanical Checks
Collimator rotation isocenter 2 mm diameter
Gantry rotation isocenter 2 mm diameter
Couch rotation isocenter 2 mm diameter
Coincidence of collimetry,gantry, couch axes with isocenter 2 mm diameter
Coincidence of radiation and mechanical isocenter 2mm diameter
Table top sag 2mm
Vertical travel of table 2mm
TG-40
Monthly
Dosimetry Tolerance
X-Ray Output Constancy 2%
Electron Output Constancy 2%
Backup monitor constancy 2%
X-ray central axis dosimetry parameter (PDD,TAR) constancy 2%
Electron central axis dosimetry parameter constancy (PDD) 2mm at theradepth
X-ray Beam Flatness constancy 2%
Electron Beam Flatness constancy 3%
X-ray and electron symmetry 3%
Mechanical Checks
Light/radiation field coincidence 2mm or 1% on a side*
Gantry/collimator angle indicators 1 degree
Wedge position 2mm (or 2% change in transmission factor)
Tray position 2mm
Applicator position 2mm
Field size indicators 2mm
Cross hair centering 2mm
Treatment couch position indicators 2mm/1deg
Latching of wedges/blocking tray Functional
Jaw symmetry 2mm
Field light intensity Functional
Monthly
Dosimetry Tolerance
X-Ray Output Constancy 2%
Electron Output Constancy 2%
Backup monitor constancy 2%
X-ray central axis dosimetry parameter (PDD,TAR) constancy 2%
Electron central axis dosimetry parameter constancy (PDD) 2mm at theradepth
X-ray Beam Flatness constancy 2%
Electron Beam Flatness constancy 3%
X-ray and electron symmetry 3%
Mechanical Checks
Light/radiation field coincidence 2mm or 1% on a side*
Gantry/collimator angle indicators 1 degree
Wedge position 2mm (or 2% change in transmission factor)
Tray position 2mm
Applicator position 2mm
Field size indicators 2mm
Cross hair centering 2mm
Treatment couch position indicators 2mm/1deg
Latching of wedges/blocking tray Functional
Jaw symmetry 2mm
Field light intensity Functional
Monday, April 14, 2008
TG-40 Daily requirements
Daily QA for Linacs
DOSIMETRY TOLERANCE
X-ray output constancy 3%
Electron output constancy*** 3%
MECHANICAL
Localizing lasers 2mm
Distance indicator (ODI) 2mm
SAFETY
Door interlock Functional
Audiovisual monitor Functional
**** All electron energies need not be checked daily, but all electron energies are to be checked at least twice weekly.
DOSIMETRY TOLERANCE
X-ray output constancy 3%
Electron output constancy*** 3%
MECHANICAL
Localizing lasers 2mm
Distance indicator (ODI) 2mm
SAFETY
Door interlock Functional
Audiovisual monitor Functional
**** All electron energies need not be checked daily, but all electron energies are to be checked at least twice weekly.
IEC
5.5.2. Safety of linac installations
The complexity of modern linacs raises concerns as to safety of operationfrom the point of view of patients and operators. The International ElectrotechnicalCommission (IEC) publishes international standards that express, asnearly as possible, an international consensus of opinion on relevant technicalsubjects; electron linacs are addressed in detail by the IEC. The IEC statementon the safety of linacs (IEC 60601-2-1, p. 13) is as follows:“The use of electron accelerators for radiotherapy purposes may exposepatients to danger if the equipment fails to deliver the required dose to thepatient, or if the equipment design does not satisfy standards of electricaland mechanical safety. The equipment may also cause danger to persons inthe vicinity if the equipment fails to contain the radiation adequately and/orif there are inadequacies in the design of the treatment room.”
The IEC document addresses three categories of safety issues —
electrical, mechanical and radiation — and establishes specific requirements
mainly for the manufacturers of linacs in the design and construction of linacs
for use in radiotherapy. It also covers some radiation safety aspects of linac
installation in customer’s treatment rooms.
The complexity of modern linacs raises concerns as to safety of operationfrom the point of view of patients and operators. The International ElectrotechnicalCommission (IEC) publishes international standards that express, asnearly as possible, an international consensus of opinion on relevant technicalsubjects; electron linacs are addressed in detail by the IEC. The IEC statementon the safety of linacs (IEC 60601-2-1, p. 13) is as follows:“The use of electron accelerators for radiotherapy purposes may exposepatients to danger if the equipment fails to deliver the required dose to thepatient, or if the equipment design does not satisfy standards of electricaland mechanical safety. The equipment may also cause danger to persons inthe vicinity if the equipment fails to contain the radiation adequately and/orif there are inadequacies in the design of the treatment room.”
The IEC document addresses three categories of safety issues —
electrical, mechanical and radiation — and establishes specific requirements
mainly for the manufacturers of linacs in the design and construction of linacs
for use in radiotherapy. It also covers some radiation safety aspects of linac
installation in customer’s treatment rooms.
ICRU recommendations
The International Commission on Radiation Units and
Measurements (ICRU, 1976) has recommended that the dose
be delivered to within 5% of the prescribed dose.
Measurements (ICRU, 1976) has recommended that the dose
be delivered to within 5% of the prescribed dose.
Thursday, April 10, 2008
8.1.3. Inverse square law (virtual source position)
In contrast to a photon beam, which has a distinct focus located at the
accelerator X ray target, an electron beam appears to originate from a point in
space that does not coincide with the scattering foil or the accelerator exit
window. The term ‘virtual source position’ was introduced to indicate the
virtual location of the electron source.
The effective source to surface distance (SSD) for electron beams
(SSDeff) is defined as the distance from the virtual source position to the point
of the nominal SSD (usually the isocentre of the linac). The inverse square law
may be used for small SSD differences from the nominal SSD to make
corrections to the absorbed dose for variations in air gaps between the patient
surface and the applicator.
There are various methods to determine the SSDeff. One commonly used
method consists of measuring the dose at various distances from the electron
applicator by varying the gap between the phantom surface and the applicator
(with gaps ranging from 0 to 15 cm). In this method, doses are measured in a
phantom at the depth of maximum dose zmax, with the phantom first in contact
with the applicator (zero gap) and then at various distances g from the
applicator. Suppose I0 is the dose with zero gap (g = 0) and Ig is the dose with
gap distance g. It follows then from the inverse square law that:
In contrast to a photon beam, which has a distinct focus located at the
accelerator X ray target, an electron beam appears to originate from a point in
space that does not coincide with the scattering foil or the accelerator exit
window. The term ‘virtual source position’ was introduced to indicate the
virtual location of the electron source.
The effective source to surface distance (SSD) for electron beams
(SSDeff) is defined as the distance from the virtual source position to the point
of the nominal SSD (usually the isocentre of the linac). The inverse square law
may be used for small SSD differences from the nominal SSD to make
corrections to the absorbed dose for variations in air gaps between the patient
surface and the applicator.
There are various methods to determine the SSDeff. One commonly used
method consists of measuring the dose at various distances from the electron
applicator by varying the gap between the phantom surface and the applicator
(with gaps ranging from 0 to 15 cm). In this method, doses are measured in a
phantom at the depth of maximum dose zmax, with the phantom first in contact
with the applicator (zero gap) and then at various distances g from the
applicator. Suppose I0 is the dose with zero gap (g = 0) and Ig is the dose with
gap distance g. It follows then from the inverse square law that:
Wednesday, April 9, 2008
Scout films
Pilot or scout films, relate CT slice position to anteriorposterior (AP) and lateral radiographic views of the patient at the time of scanning.
They are obtained by keeping the X-ray source in a fixed position and moving the patient (translational motion) through the stationary slit beam. The result is a high defininition radiograph that is divergent on the transverse axis but non-divergent on the longitudinal axis.
Target position relative to the bony anatomy on the simulator radiographs may then be determined through comparison with the CT scout or pilot films, keeping in mind the different magnifications between the simulator and scout films.
They are obtained by keeping the X-ray source in a fixed position and moving the patient (translational motion) through the stationary slit beam. The result is a high defininition radiograph that is divergent on the transverse axis but non-divergent on the longitudinal axis.
Target position relative to the bony anatomy on the simulator radiographs may then be determined through comparison with the CT scout or pilot films, keeping in mind the different magnifications between the simulator and scout films.
Saturday, April 5, 2008
Electron beam and virtual SSD
Electron beams are used with external collimating devices known as
cones or applicators that reduce the spread of the electron beam in the air. The
design of these cones is dependent on the manufacturer and affects the
dosimetric properties of the beam.
Electron shielding for irregular fields may be accomplished with the use
of thin lead or low melting point alloy inserts. These shielding inserts can have
significant effects upon the electron beam dosimetry (especially the PDD and
output), and these effects may be modelled by the TPS.
The design of the linac head may be important for electron dosimetry,
especially for Monte Carlo type calculations. In these conditions particular
attention is paid to the scattering foil. The effective or virtual SSD will appear
to be shorter than the nominal SSD, and may be taken into consideration by
the TPS.
Podgorsak
cones or applicators that reduce the spread of the electron beam in the air. The
design of these cones is dependent on the manufacturer and affects the
dosimetric properties of the beam.
Electron shielding for irregular fields may be accomplished with the use
of thin lead or low melting point alloy inserts. These shielding inserts can have
significant effects upon the electron beam dosimetry (especially the PDD and
output), and these effects may be modelled by the TPS.
The design of the linac head may be important for electron dosimetry,
especially for Monte Carlo type calculations. In these conditions particular
attention is paid to the scattering foil. The effective or virtual SSD will appear
to be shorter than the nominal SSD, and may be taken into consideration by
the TPS.
Podgorsak
Styrofoam insert in an HDR calibration well chamber unit
I have seen in the past a question on the ABR Oral exams about what the styrofoam insert in an HDR calibration well chamber is for. FOr almost a year I've searched for a definite anser to this problem with no luck. I asked a friend who used to work at the RPC, and he said he thought it had to do with the strength of a 10 Ci Ir192 source being strong enought to rasie the temperature in the chamger and tus change its response. --But that was just his guess he told me.Finally, I found a real-live reference that gives a solid answer.AAPM Task Group-41 (page 77) states:"The well chambers designed for HDR sourcesexhibit a heat dependency, e.g., the 370.-GBq (10.0-Ci) sources canincrease the air temperature in the collection well if they remain therefor several minutes.72 Styrofoam thermal absorbers can be positionedaround the source holder to alleviate this problem."I believe this answer along with a mention of TG41 would suffice as an answer for an ABR examiner.
13.1 Field Blocks
A. Block Thickness
Shielding blocks are primarily made of lead. A thickness of 4-5 HVL’s rings the initial
beam transmission to 5%.
1/2n = 0.05
n log 2 = log (1/0.05)
n = 4.32 HVL’s to reduce a beam to 5% transmission
A. Block Thickness
Shielding blocks are primarily made of lead. A thickness of 4-5 HVL’s rings the initial
beam transmission to 5%.
1/2n = 0.05
n log 2 = log (1/0.05)
n = 4.32 HVL’s to reduce a beam to 5% transmission
Corrections for Contour Irregularities (Khan Ch 12)
12.4 Corrections for Contour Irregularities
The surface of the patient may be irregular and the beam may be oblique with respect to the patient. In order to correct these conditions the following methods are proposed:
A. Effective SSD Method
If the SSD is large (>80 cm) then the PDD will not change rapidly with respect to changing SSD in an irregular patient contour.
PDD2 = PDD1 • [SSD + dmax/SSD + Δd + dmax]2
B. TAR or TMR ratio method (overestimates the dose for all energies)
TAR and TMR do not depend on SSD and are only a function of depth and field size.
Correction factor CF = TAR(d, ra)/ TAR(d + Δd, ra)
PDD2 = PDD1 • CF
C. Isodose shift method
While the 2 preceding methods are good for individual spots, an entire chart may be corrected by superimposing a grid that shifts all of the PDD lines with respect with the missing tissue.
The surface of the patient may be irregular and the beam may be oblique with respect to the patient. In order to correct these conditions the following methods are proposed:
A. Effective SSD Method
If the SSD is large (>80 cm) then the PDD will not change rapidly with respect to changing SSD in an irregular patient contour.
PDD2 = PDD1 • [SSD + dmax/SSD + Δd + dmax]2
B. TAR or TMR ratio method (overestimates the dose for all energies)
TAR and TMR do not depend on SSD and are only a function of depth and field size.
Correction factor CF = TAR(d, ra)/ TAR(d + Δd, ra)
PDD2 = PDD1 • CF
C. Isodose shift method
While the 2 preceding methods are good for individual spots, an entire chart may be corrected by superimposing a grid that shifts all of the PDD lines with respect with the missing tissue.
Monday, March 31, 2008
Linac based stereotactic treatment procedures
Linac based stereotactic treatment procedures are:
a) single plane transverse rotation (Gamma Knife)
b) multiple noncoplanar converging arcs
c) single arc dynamic rotation
a) single plane transverse rotation (Gamma Knife)
b) multiple noncoplanar converging arcs
c) single arc dynamic rotation
Cylindrical tertiary collimators (cones)
Use of cylindrical tertiary collimators in stereotactic radiotherapy
a) provides more precise collimation
b) gives rise to a spherical dose distribution surrounding the target volume
c) gives sharper dose fall off compared to secondary collimators
a) provides more precise collimation
b) gives rise to a spherical dose distribution surrounding the target volume
c) gives sharper dose fall off compared to secondary collimators
Gamma Knife
Gamma Knife radiosurgery
201 Co-60 sources distributed isotropically in a hemisphere
Collimator diameters vary between 4 and 18 mm.
201 Co-60 sources distributed isotropically in a hemisphere
Collimator diameters vary between 4 and 18 mm.
Stereotactic with 6 MV
For a 6MV photon beam used for stereotactic radiosurgery, the minimum beam size diamter (in mm) required for electronic equilibrium is about 30 mm.
Dosimetry of small stereotactic fields
Dosimetry of small stereotactic fields involves:
The problem of lack of charged particle equilibrium be given special consideration and also that TMR and off axis ratios be measured for small fields.
With Co-60 beam treatment, one requires about 5 mm of tissue material for charged particle equilibrium. So a field size of less than about 1 cm will give reduced output due to lack of CPE. Similarly for a 6MV beam, the minimum field size for CPE is about 3 cm; though in practice, the condition may hold slightly smaller field sizes as well since the buildup curve is very steep.
So dosimetric functions must be measured for very small field sizes, for treatment purposes. Also such small fields require microchambers for measurement.
A Farmer type chamber has typical dimensions of about 16 to 18 mm diameter by about 22 mm length, which is not suitable for the dosimetry of small fields.
The solution is a PTW pinpoint chamber (SEE OR INSERT DIAGRAM)
The problem of lack of charged particle equilibrium be given special consideration and also that TMR and off axis ratios be measured for small fields.
With Co-60 beam treatment, one requires about 5 mm of tissue material for charged particle equilibrium. So a field size of less than about 1 cm will give reduced output due to lack of CPE. Similarly for a 6MV beam, the minimum field size for CPE is about 3 cm; though in practice, the condition may hold slightly smaller field sizes as well since the buildup curve is very steep.
So dosimetric functions must be measured for very small field sizes, for treatment purposes. Also such small fields require microchambers for measurement.
A Farmer type chamber has typical dimensions of about 16 to 18 mm diameter by about 22 mm length, which is not suitable for the dosimetry of small fields.
The solution is a PTW pinpoint chamber (SEE OR INSERT DIAGRAM)
Sunday, March 30, 2008
Adjacent photon fields abutting at depth of dose specification
In a treatment involving adjacent photon fields abutting at depth of dose specification, the overlap region (below the junction gets overdosed and the region above the junction gets overdosed).
In order to make the dose more uniform:
1) The gap on the surface can be increased or decreased in successive fractions
2) the two fields can be angled in opposite directions to make the beam edges parallel at the junction
3) The two fields can be treated by half fields (using half beam blocks or asymmetric jaws), the central axes becoming the matching edges of the two adjacent fields
4) A matching pair of POP fields can be used.
FOLLOWUP
Adjacent fields abut at the depth where the dose is required to be uniform (over the full length of the two abutting fields).
This however, creates high and low dose regions around the abutting junction. The high dose region is created by the adjacent fields diverging into each other. In order to accept this plan for treatment, the low end high dose regions must be free of tumor and organ at risk, respectively.
The "hot" and "cold" regions can be avoided by the various methods mentioned in this question, namely making the abutting field borders parallel to each other.
In order to make the dose more uniform:
1) The gap on the surface can be increased or decreased in successive fractions
2) the two fields can be angled in opposite directions to make the beam edges parallel at the junction
3) The two fields can be treated by half fields (using half beam blocks or asymmetric jaws), the central axes becoming the matching edges of the two adjacent fields
4) A matching pair of POP fields can be used.
FOLLOWUP
Adjacent fields abut at the depth where the dose is required to be uniform (over the full length of the two abutting fields).
This however, creates high and low dose regions around the abutting junction. The high dose region is created by the adjacent fields diverging into each other. In order to accept this plan for treatment, the low end high dose regions must be free of tumor and organ at risk, respectively.
The "hot" and "cold" regions can be avoided by the various methods mentioned in this question, namely making the abutting field borders parallel to each other.
CSI treatments
In CSI treatments, the cranial field on the abutting side can be made nondiverging by using a half beam block. The size of the abutting spinal field is 40cm x 6 cm. SSD is 100 cm.
If Pheta is the angle of rotation of the collimator required for the abutting of the cranial and spinal fields, Tan Pheta is given by
In this particular example, the abutting border has no divergence. So collimator rotation to match the divergence of the spinal field would be enough to match the two orthogonal fiedlds.
The inferior border of the brain field in this case, will not diverge into the abutting border of the spine field.
If Pheta is the angle of rotation of the collimator required for the abutting of the cranial and spinal fields, Tan Pheta is given by
In this particular example, the abutting border has no divergence. So collimator rotation to match the divergence of the spinal field would be enough to match the two orthogonal fiedlds.
The inferior border of the brain field in this case, will not diverge into the abutting border of the spine field.
Friday, March 28, 2008
TBI Treatment
In TBI treatment, the
1) prescribed dose is about 1200 cGy, generally delivered using POP of fields
2) dose is delivered as a single fraction
3) dose is delivered as multiple fractions, 10 to 15 cGy per fraction
High dose TBI is delivered in a single fraction or in a small number of fractions of 200 cGy/fraction, 6 fractions for a total 1200 cGy.
Low dose TBI is delivered in 10 to 15 fractions of 10 to 15 cGy per fraction. For further study see Podgorsak.
1) prescribed dose is about 1200 cGy, generally delivered using POP of fields
2) dose is delivered as a single fraction
3) dose is delivered as multiple fractions, 10 to 15 cGy per fraction
High dose TBI is delivered in a single fraction or in a small number of fractions of 200 cGy/fraction, 6 fractions for a total 1200 cGy.
Low dose TBI is delivered in 10 to 15 fractions of 10 to 15 cGy per fraction. For further study see Podgorsak.
Total body irradiation
Total body irradiation involves:
Irradiation of the whole body of the patient
Delivery of dose in single or multiple fractions
Use of Co-60 beams or accelerator photon beams
NOTE +/- 10% accuracy for both TBI and TSET
Irradiation of the whole body of the patient
Delivery of dose in single or multiple fractions
Use of Co-60 beams or accelerator photon beams
NOTE +/- 10% accuracy for both TBI and TSET
Tuesday, March 25, 2008
Gamma Analysis
Gamma Analysis
Although the use of the two factors provides the
independent evaluation of a dose difference and misalignment,
the gamma offers a composite analysis
with the two variables collapsed into one parameter
(Harms et al. 1998; Low et al. 1998c; Dupuydt,Van
Esch, and Huyskens 2002). The gamma is defined as
the square root of a linear quadratic addition of the
two factors, while they are provided in relative magnitude
to their acceptance criteria (CDTA and CDD), as
shown in equation (4.1).
Parallel Plate Chamber Design (Govinda Rajan)
PP chamber design is very simple.
1) Take two circular perspex plates of about 1 mm thickness and about 4or 5 cm in diameter.
2)Coat the inside of one disk with graphite(aquadaq).
3) Take the other disk and keep a thin ring of less than or equal to one mm thickness and 1 cm in diameter and keep it concentric with the disk centre and now coat with graphite. You now have the central collecting electrode and the outer guard with a smallinsulating gap in between.
4) Take a perspex ring of anout 5 cm diameterand about 2 mm thickness and fix the two disks on the ring with the graphite coated portion on the inside. you now have the high tension electrode, the central electrode and guard.
5)Now you have to take the connections. HT wall is the entry window. If you want dose at smaller depths you must make this wall thinner. If you want the cavity thickness to be smaller than 2 mm you must use a ring of smaller thickness as spacer between the two electrodes. If you want a larger electrode you must use a larger inner ring before coating withgraphite.If you need a larger guard, the diameter of the perspex disk chosen must be larger.
If you see the protocol. you will find PP chambers of different sizes because of changes in the parameters thatwe discussed. How the characteristics will depend on these dimensions are given in the protocol.If you can make the spacer ring whose thickess can be varied, you have an extrapolation chamber. If you make the wall materials interchangeable you can study wall effects.
You need some expertise to attach wires to the central electrode,guard and HT wall and take out and connect to cable.Hope this gives you some idea of PP chamber
Govinda Rajan.K.N.Govinda Rajan.K.N.
1) Take two circular perspex plates of about 1 mm thickness and about 4or 5 cm in diameter.
2)Coat the inside of one disk with graphite(aquadaq).
3) Take the other disk and keep a thin ring of less than or equal to one mm thickness and 1 cm in diameter and keep it concentric with the disk centre and now coat with graphite. You now have the central collecting electrode and the outer guard with a smallinsulating gap in between.
4) Take a perspex ring of anout 5 cm diameterand about 2 mm thickness and fix the two disks on the ring with the graphite coated portion on the inside. you now have the high tension electrode, the central electrode and guard.
5)Now you have to take the connections. HT wall is the entry window. If you want dose at smaller depths you must make this wall thinner. If you want the cavity thickness to be smaller than 2 mm you must use a ring of smaller thickness as spacer between the two electrodes. If you want a larger electrode you must use a larger inner ring before coating withgraphite.If you need a larger guard, the diameter of the perspex disk chosen must be larger.
If you see the protocol. you will find PP chambers of different sizes because of changes in the parameters thatwe discussed. How the characteristics will depend on these dimensions are given in the protocol.If you can make the spacer ring whose thickess can be varied, you have an extrapolation chamber. If you make the wall materials interchangeable you can study wall effects.
You need some expertise to attach wires to the central electrode,guard and HT wall and take out and connect to cable.Hope this gives you some idea of PP chamber
Govinda Rajan.K.N.Govinda Rajan.K.N.
Shielding notes
1. Shielding Design Goal
The shielding design goals (P) are 0.02 mSv/week (1mSv/yr) for uncontrolled areas and 0.1 mSv/week (5mSv/yr) for controlled areas.
The maximum dose equivalent in any one hour (TADR) is 0.02 mSv (20µSv).
SOURCE: Shielding Calculation Report
Varian iX
The shielding design goals (P) are 0.02 mSv/week (1mSv/yr) for uncontrolled areas and 0.1 mSv/week (5mSv/yr) for controlled areas.
The maximum dose equivalent in any one hour (TADR) is 0.02 mSv (20µSv).
SOURCE: Shielding Calculation Report
Varian iX
Monday, March 24, 2008
TSEI Total skin electron irradiation
Total skin electron irradiation involves:
1) treating large areas of skin and it's underlying layers
2) use of low energy (about 3 to 6 MeV) electron beams
3) measurement of beam uniformity and output for the treatment geometry
4) an accuracy of about +/ 10%
1) treating large areas of skin and it's underlying layers
2) use of low energy (about 3 to 6 MeV) electron beams
3) measurement of beam uniformity and output for the treatment geometry
4) an accuracy of about +/ 10%
Electron arc therapy
Electron arc therapy
1) it is used to treat superificial layers of skin that are not curved
2) it is used to create superficial layers that have curvature as well, ex the chest wall
3) it requires proper dosimetric procedures to be utilized
1) it is used to treat superificial layers of skin that are not curved
2) it is used to create superficial layers that have curvature as well, ex the chest wall
3) it requires proper dosimetric procedures to be utilized
Clarksons method
Using Clarksons method:
1) dose at any point in a patient can be computed for an irregular field
2) scatter and primary components of dose at any point in a patient can be evaluated.
1) dose at any point in a patient can be computed for an irregular field
2) scatter and primary components of dose at any point in a patient can be evaluated.
Blocking
When a part of the open beam is blocked for reducing dose to an organ at risk,
Primary dose in the patient under the unblocked region remains almost the same.
Moderate blocking does not affect the effective primary reaching the phantom. Scatter dose under the unblocked region decreases since the shadow region woudl scatter very little radiation.
The dose under the block comprises the primary leakage radiaton and scatter going into the shadow region from unblocked regions.
Primary dose in the patient under the unblocked region remains almost the same.
Moderate blocking does not affect the effective primary reaching the phantom. Scatter dose under the unblocked region decreases since the shadow region woudl scatter very little radiation.
The dose under the block comprises the primary leakage radiaton and scatter going into the shadow region from unblocked regions.
LDR, MDR, HDR
DOSE RATE Dose Rate at the dose specification point
Low dose rate (LDR) 0.4-2 Gy/h
Medium Dose Rate (MDR) 2-12 Gy/h
High Dose Rate (HDR) >12 Gy/h
Low dose rate (LDR) 0.4-2 Gy/h
Medium Dose Rate (MDR) 2-12 Gy/h
High Dose Rate (HDR) >12 Gy/h
Sunday, March 23, 2008
CT Numbers
Because CT numbers bear a linear relationship with the attenuation coefficients it is possible to infer electron density (electrons cm-3). Although CT numbers can be correlated with electron density, the relationship is not linear in the entire range of tissue densities. The nonlinearity is caused by change in atomic number of tissues, which affects the proportion of beam attenuation by Compton versus photoelectric interactions. Khan (see Figure on pg 233)
To ensure accurate dose calculation, the CT numbers must be converted
to electron densities and scattering powers. The conversion of CT numbers to
electron density and scattering power is usually performed with a user defined
look-up table, which in turn is generated using a water equivalent circular
phantom containing various inserts of known densities simulating normal body
tissues such as bone and lung. Pdgorsak.
To ensure accurate dose calculation, the CT numbers must be converted
to electron densities and scattering powers. The conversion of CT numbers to
electron density and scattering power is usually performed with a user defined
look-up table, which in turn is generated using a water equivalent circular
phantom containing various inserts of known densities simulating normal body
tissues such as bone and lung. Pdgorsak.
DRR
Digitally reconstructed radiographs
DRRs are produced by tracing ray lines from a virtual source position
through the CT data of the patient to a virtual film plane. The sum of the
attenuation coefficients along any one ray line gives a quantity analogous to
optical density (OD) on a radiographic film. If the sums along all ray lines from
a single virtual source position are then displayed on to their appropriate
positions on the virtual film plane, the result is a synthetic radiographic image
based wholly on the 3-D CT data set that can be used for treatment planning.
Figure 7.10 provides an example of a typical DRR.
Pdgorsak
DRRs are produced by tracing ray lines from a virtual source position
through the CT data of the patient to a virtual film plane. The sum of the
attenuation coefficients along any one ray line gives a quantity analogous to
optical density (OD) on a radiographic film. If the sums along all ray lines from
a single virtual source position are then displayed on to their appropriate
positions on the virtual film plane, the result is a synthetic radiographic image
based wholly on the 3-D CT data set that can be used for treatment planning.
Figure 7.10 provides an example of a typical DRR.
Pdgorsak
Friday, March 21, 2008
Depth shifting ionization curves for gradient effects (TG-51)
Figure 1: Effect of shifting depth-ionization data measured with cylindrical chambers upstream by 0.6 for photon beams (panel a) and 0.5 for electron beams (panel b) (with = 1.0 cm). The raw data are shown by curve I (long dashes) in both cases and the shifted data, which are taken as the depth-ionization curve, are shown by curve II (solid line). The value of the % ionization at point A (10 cm depth) in the photon beam gives and the depth at point B (solid curve, 50% ionization) in the electron beam gives from which can be determined (see section VIII.C). For the photon beams, curve II is effectively the percentage depth-dose curve. For the electron beams, curve II must be further corrected (see section X.D) to obtain the percentage depth-dose curve shown (short dashes - but this is not needed for application of the protocol).
Thursday, March 20, 2008
Craniospinal irradiation (CSI)
While treating a patient for craniospinal irradiation (CSI), two orthogonal fields must abut at the desired depth in the patient. If both fields are divergent, for the abutting of lateral and spinal fields:
Both collimator and couch must be turned through the angle of divergence of abutting spinal and cranial fields respectively.
See a figure in the Dosimetry Review Book.
The figure shows the rotation of the collimator by an angle to make the inferior border of the lateral field tangent to the P/A spinal field for the abutting of two orthogonal fields.
Both collimator and couch must be turned through the angle of divergence of abutting spinal and cranial fields respectively.
See a figure in the Dosimetry Review Book.
The figure shows the rotation of the collimator by an angle to make the inferior border of the lateral field tangent to the P/A spinal field for the abutting of two orthogonal fields.
Blocking and collimator scatter factor (Sc)
Example
In a radiation therapy treatment, using a Co-60 beam, the open field, 12 cmx12 cm requires some blocking at the periphery to protect normal structures around PTV. The blocked field was 8cm x 8 cm in size so:
Sc of blocked field = Sc of unblocked field
BECAUSE, blocks are below the area of collimator scatter.
In a radiation therapy treatment, using a Co-60 beam, the open field, 12 cmx12 cm requires some blocking at the periphery to protect normal structures around PTV. The blocked field was 8cm x 8 cm in size so:
Sc of blocked field = Sc of unblocked field
BECAUSE, blocks are below the area of collimator scatter.
Partial blocking
When a blocked field is not a significant fraction of an unblocked or open field;
1) the effective primary reaching the patient is altered to a negligible amount
2) the collimator scatter factor Sc remains the same as that of the unblocked field
Blocking does change scatter volume and hence the phantom scatter factor (Sp) changes.
The output factor also changes relative to the unblocked field.
1) the effective primary reaching the patient is altered to a negligible amount
2) the collimator scatter factor Sc remains the same as that of the unblocked field
Blocking does change scatter volume and hence the phantom scatter factor (Sp) changes.
The output factor also changes relative to the unblocked field.
Wednesday, March 19, 2008
SAR (scatter air ratio)
Scatter Air Ratio (SAR) is obtained by subtracting zero field TAR from TAR of the given field. It is generally required for the dosimetry of irregular fields.
It dose depend on field size since scatter volume increases with an increase in field size.
It dose depend on field size since scatter volume increases with an increase in field size.
Photon beam dose profile
The photon beam dose profile at depth of clinical interest is fairly uniform across the field size for Co-60 and linac beams.
It is nonuniform across field size for linac photon beams due to flattening filter effects.
It is nonuniform across field size for linac photon beams due to flattening filter effects.
Field blocking
Field blocking
1) reduces scatter volume
2) reduces scatter dose at the depth of interest
3) leaves the primary dose reaching the phantom along the beam central axis relatively unaffected (assuming that blocking is neither along central axis nor a larger fraction of open field)
1) reduces scatter volume
2) reduces scatter dose at the depth of interest
3) leaves the primary dose reaching the phantom along the beam central axis relatively unaffected (assuming that blocking is neither along central axis nor a larger fraction of open field)
Scatter dose
The scatter dose at any point in a phantom can be determined
1) by the evaluation of the scatter air ratio (SAR)
2) by the evaluation of the scatter maximum ratio
1) by the evaluation of the scatter air ratio (SAR)
2) by the evaluation of the scatter maximum ratio
Equivalent square field of an irregular field
The equivalent square field of any irregular field for any dose function can be determined by
CLARKSONS sector integration method
CLARKSONS sector integration method
Scatter Air Ratio
Scatter Air Ratio gives the scatter component of TAR
It is the difference between TAR and zero field TAR
The concept aids in separating the scatter and primary dose at a point in the patient
It is the difference between TAR and zero field TAR
The concept aids in separating the scatter and primary dose at a point in the patient
Independent jaw movement
Independent jaws in a linac aid in
1) Field blocking
2) Field splitting
3) Field matching
4) creating dynamic wedge fields
1) Field blocking
2) Field splitting
3) Field matching
4) creating dynamic wedge fields
Lead Thickness and electrons
The thickness of the lead cutout you would use with a 10 MeV electron is about:
5mm
Electron absorption is about 2 MeV/CM of unit density material (water) and 10 times less for lead so 2MeV/mm for lead.
5mm
Electron absorption is about 2 MeV/CM of unit density material (water) and 10 times less for lead so 2MeV/mm for lead.
Stereotactic Procedure
Image Transfer
Fixer (to remove unneeded segments below and above)
Localize the CT (selecting the rods) first before fusing. Fuse first CT axial to MR axial, then CT axial to MR coronal.
Save PTV in terms of it's contour as PTVaxial and PTV coronal to avoid confusion.
Create body contour and trim.
Plan by manipulating percentages to the target of interest and organs at risk
Fixer (to remove unneeded segments below and above)
Localize the CT (selecting the rods) first before fusing. Fuse first CT axial to MR axial, then CT axial to MR coronal.
Save PTV in terms of it's contour as PTVaxial and PTV coronal to avoid confusion.
Create body contour and trim.
Plan by manipulating percentages to the target of interest and organs at risk
Monday, March 17, 2008
Wedged treatment pairs
When using a pair of wedged treatment fields, the wedge angle is much smaller than the ideal one. The anterior region of overlap will lead to
a high dose region.
Conversely, when using a pair of wedged treatment fields and the wedge angle is much larger than the ideal one, the anterior region of overlap will lead to a low dose region
a high dose region.
Conversely, when using a pair of wedged treatment fields and the wedge angle is much larger than the ideal one, the anterior region of overlap will lead to a low dose region
Hinge angle and wedge angle PART II
In a treatment involving a pair of wedged fields, the hinge angle is 90 degrees. What wedge angle must be chosen to get uniform dose to the region of overlap?
90 - FI/2, where FI is the hinge angle.
90 - FI/2, where FI is the hinge angle.
Wedge Angle and Hinge Angle
Making dose in an overlap region uniform:
This can occur by wedging the two fields, choosing the proper wedge angle that would make the isodose lines parallel to one another in the overlap region.
Finally this can be performed by correcting for surface obliquity if or after wedging the fields.
NOTE:
For a given hinge angle, there is a wedge angle for a wedge pair of fields that would make the dose uniform in the beam overlap region.
This assumes that the incident surface is normal to the central axis. If the patient surface is sloping, either compensator should be made use of or the wedge angle, derived from hinge angle should be adjusted to take into account the hinge angle.
This can occur by wedging the two fields, choosing the proper wedge angle that would make the isodose lines parallel to one another in the overlap region.
Finally this can be performed by correcting for surface obliquity if or after wedging the fields.
NOTE:
For a given hinge angle, there is a wedge angle for a wedge pair of fields that would make the dose uniform in the beam overlap region.
This assumes that the incident surface is normal to the central axis. If the patient surface is sloping, either compensator should be made use of or the wedge angle, derived from hinge angle should be adjusted to take into account the hinge angle.
Sunday, March 16, 2008
Bolus
The use of bolus in electron beam therapy is obviously used to enhance the skin dose, but also can be used for depth compensation as well (for example, if there is great variation in the chest wall thickness).
Lead vs Cerrobend
The thickness of lead required for making a field block for a clinical photon beam is about 8 cm. If Cerrobend material is used, the required thickness would be:
ABOUT 20% MORE. This is because cerrobend is 20% less dense than lead.
ABOUT 20% MORE. This is because cerrobend is 20% less dense than lead.
Blocks vs MLC
Disadvantages in using custom blocks are:
Small errors in block position can adversely affetct treatment and fabrication, alignment, etc take alot of time.
This keys in to the advantages of MLC, which also include the fact that treatment delivery is faster and they are equally convenient for any angled field.
Small errors in block position can adversely affetct treatment and fabrication, alignment, etc take alot of time.
This keys in to the advantages of MLC, which also include the fact that treatment delivery is faster and they are equally convenient for any angled field.
Thickness of blocks
The thickness of blocks used for field blocking (in HVL) is about 5 HVL's.
To reduce the dose in the shielding region to less than 5% of the open field dose, the shield thickness must be greater than 5HVL's
To reduce the dose in the shielding region to less than 5% of the open field dose, the shield thickness must be greater than 5HVL's
Primary transmission for a block
The primary transmission for a block used to shield a portion of the field is usually about
3% to 5%
The dose under a block in a patient will be MORE than 3% to 5% and depends on how much patient scatter the shadow region of the patient receives.
3% to 5%
The dose under a block in a patient will be MORE than 3% to 5% and depends on how much patient scatter the shadow region of the patient receives.
Errors in custom field blocking
Errors in custom field blocking can arise due to
1) incorrect source to film distance (SFD)
2) incorrect source to tray distance
3) incorrect positioning of the block on the tray.
Incorrect SFD will change the size of the block fabricated
Incorrect source to tray distance or incorrect repositioning of the block or tray will influence the region blocked in patient.
1) incorrect source to film distance (SFD)
2) incorrect source to tray distance
3) incorrect positioning of the block on the tray.
Incorrect SFD will change the size of the block fabricated
Incorrect source to tray distance or incorrect repositioning of the block or tray will influence the region blocked in patient.
Physical Penumbra
The physical penumbra width is defined as the lateral distance between two specified isodose curves at a specified depth (e.g., lateral distance between 90% and 20% isodose lines at the depth of dmax).
Geometric penumbra
Geometric penumbra, which exists both inside and outside the geometrical boundaries of the beam depends on:
Source size
Distance from the source and
Source to diaphragm distance.
Source size
Distance from the source and
Source to diaphragm distance.
"HORNS"
Linac x-ray beams exhibit areas of high dose known as horns, near the surface in the periphery in the field.
These horns are created by the flattening filter, which is usually designed to overcompensate near the surface in order to obtain flat isodose curves at great depths.
These horns are created by the flattening filter, which is usually designed to overcompensate near the surface in order to obtain flat isodose curves at great depths.
Saturday, March 15, 2008
Monday, March 10, 2008
Imaging book
Radiology Review: Radiologic Physics (Paperback)by Edward L. Nickoloff (Author), Naveed Ahmad (Author)Price: $56.95
Will take a look for this at med school bookstore. Someone on yahoo group said reviews recommended it over Huda.
Will take a look for this at med school bookstore. Someone on yahoo group said reviews recommended it over Huda.
Custom blocks
Custom blocks used for shielding critical structures in the treatment of MV photon beams must be diverging.
THEY ARE NOT KEPT TO CLOSE TO THE PATIENT SKIN BECAUSE THAT WOULD INCREASE THE ELECTRON CONTAMINATION REACHING THE PATIENT
However, also consider this: Another reason not to keep it close to the patient is that if the block is closer to the patient, the larger size and weight of the block needes to shield the same volume will be a disadvantage. This occurs due to the beam divergence.
On the other hand, the transmission penumbra would be larger for a larger block to skin distance. A block to skin distance of about 15 to 20 cm is a good compromise for positioning custom blocks in patient treatment.
THEY ARE NOT KEPT TO CLOSE TO THE PATIENT SKIN BECAUSE THAT WOULD INCREASE THE ELECTRON CONTAMINATION REACHING THE PATIENT
However, also consider this: Another reason not to keep it close to the patient is that if the block is closer to the patient, the larger size and weight of the block needes to shield the same volume will be a disadvantage. This occurs due to the beam divergence.
On the other hand, the transmission penumbra would be larger for a larger block to skin distance. A block to skin distance of about 15 to 20 cm is a good compromise for positioning custom blocks in patient treatment.
Field weighting
Field weighting
1) gives the relative contribution of beams to dose at target center or dmax point
2) improves dose uniformity across the target
3) reduces dose to normal tissues or critical structures
Field weighting is used:
When contribution from any of the fields needs to be reduced or increased with respect to another field.
1) gives the relative contribution of beams to dose at target center or dmax point
2) improves dose uniformity across the target
3) reduces dose to normal tissues or critical structures
Field weighting is used:
When contribution from any of the fields needs to be reduced or increased with respect to another field.
Dose rate and extended SSD
In the previous post, we have a patient with a 6 MV photon beam at an extended SSD of 125 cm. The calibration dose rate is given at Dcal (1.5, 10x10, 100) by 0.993 cGy/MU.
Calculate the reference dose rate in cGy/MU for the new SSD:
ANSWER
Obviously we know the dose rate in cGy/MU will decrease with a change to the extended SSD distance. The extent is given by
Calib dose rate x (100+dmax)/(extended SSD+ dmax) = 0.993 (100+1.5)/(125+1.5)
Calculate the reference dose rate in cGy/MU for the new SSD:
ANSWER
Obviously we know the dose rate in cGy/MU will decrease with a change to the extended SSD distance. The extent is given by
Calib dose rate x (100+dmax)/(extended SSD+ dmax) = 0.993 (100+1.5)/(125+1.5)
Treatment at extended SSD
A patient is to be treated with a 6MV photon beam for a field size of 12cmx12cm on the patient surface, at an extended SSD of 125 cm. (PLEASE SKETCH DIAGRAM). The depth of target center is 9 cm. In this case (compared to a standard SSD treatment):
1) The reference dose rate (or dose/MU) at the input port (dmax position) decreases relative to a shorter SSD
2) the PDD (9, 12x12, 125) increases due to the increased SSD
3) the PDD correction is approximately given by Mayneords F Factor
1) The reference dose rate (or dose/MU) at the input port (dmax position) decreases relative to a shorter SSD
2) the PDD (9, 12x12, 125) increases due to the increased SSD
3) the PDD correction is approximately given by Mayneords F Factor
Extended SSD treatments
Extended SSD treatments are used for:
1) Total Body irradiation
2) Large mantle fields
1) Total Body irradiation
2) Large mantle fields
Clarksons method
Clarksons method is useful for dose calculation of irregular fields. (i.e. Mantle cases, per example in Khan)
Entrance dose and exit dose
Entrance dose is defined as the dose at dmax for the incident field.
The exit dose is defined as the dose at dmax for the exiting field.
The exit dose is defined as the dose at dmax for the exiting field.
Collimator scatter
The collimator scatter (Sc) for accelerator photon beams can be measured in air with the Farmer chamber with a cap of appropriate build-up thickness.
Increase in field size
With an increase in field size, the (effective) primary radiation incident on a patient increases.
Minimizing gap between two adjacent fields
To minimize the gap between two adjacent fields:
A half beam block can be used.
A half beam block can be used.
Physical Penumbra
Physical penumbra depends on
Geometric penumbra
Collimator transmission
Lateral Photon Scatter (in the patient)
Lateral Electron Transport (in the patient)
Geometric penumbra
Collimator transmission
Lateral Photon Scatter (in the patient)
Lateral Electron Transport (in the patient)
Beam quality and Peak Scatter Factor (PSF)
Beam quality and PSF for a 15cmx15 cm field
a) Co-60 beam - 1.05
b) 6 MV- 1.015
c) 18 MV- 1.008
a) Co-60 beam - 1.05
b) 6 MV- 1.015
c) 18 MV- 1.008
Beam Quality and PDD (10x10, 10 deep)
Beam Quality and PDD (10x10, 10 deept), SSD=80 cm for Co-60 and 100 cm for accelerator photon beams.
BEAM PDD
Co-60 55.6
4 MV 64.8
10 MV 73
BEAM PDD
Co-60 55.6
4 MV 64.8
10 MV 73
Overall Uncertainty in Dose Delivery
Per Khan Ed 3, Table 17.4, pg 431.
Step Uncertainty (%)
Ion chamber calibration 1.6
Calibration procedure 2.0
Dose calc parameters and methods 3.0
Effective depth 2.0
SSD 2.0
Wedges 2.0
Blocking Trays 2.0
CUMULATIVE 5.6
Step Uncertainty (%)
Ion chamber calibration 1.6
Calibration procedure 2.0
Dose calc parameters and methods 3.0
Effective depth 2.0
SSD 2.0
Wedges 2.0
Blocking Trays 2.0
CUMULATIVE 5.6
Beam quality and dmax for a 10x10 field
Beam quality and depth of dmax for a 10x10 cm field
Kilovoltage- 0 cm
Co-60 - 0.5 cm
10 MV- 2.5 cm
25 MV- 3.5 cm
Kilovoltage- 0 cm
Co-60 - 0.5 cm
10 MV- 2.5 cm
25 MV- 3.5 cm
Surface dose for clinical electron beams
The surface dose for clinical electron beams is about 75% to 95%
Surface dose for electron beams is much larger compared to photon beams and can be measured with an extrapolation chamber.
Surface dose for electron beams is much larger compared to photon beams and can be measured with an extrapolation chamber.
Electron energies (in MeV) and their ranges in water
Range in water for various electron energies
6 MeV ------- 3 cm
8 Mev ------- 4 cm
10 Mev------ 4.9 cm
20 MeV----- 9.2 cm
The range in cm is approximately related to the energy in MeV/2
6 MeV ------- 3 cm
8 Mev ------- 4 cm
10 Mev------ 4.9 cm
20 MeV----- 9.2 cm
The range in cm is approximately related to the energy in MeV/2
Sunday, March 9, 2008
Thursday, March 6, 2008
Mean Energy of the electron beam (Eo)
The mean energy of the electron beam, Eo, can be determined, fairly accurately, from the parameter
2.33 R50
2.33 R50
Corrections for ion chamber measurements
For accurate dosimetry, chamber measurements must be corrected for
a) saturation
b) polarity
c) temperature/pressure
a) saturation
b) polarity
c) temperature/pressure
Electron energy specified by the machine manufacturer (II)
The electron energy specified by the machine manufacturer refers to:
The most probable energy of the electrons. The manufacturer usually measures the practical range of the electron beam, which is related to the most probable energy of the electron beam.
The most probable energy of the electrons. The manufacturer usually measures the practical range of the electron beam, which is related to the most probable energy of the electron beam.
Electron energy specified by the manufacturer
The electron energy beam incident on a patient can be characterized by
a) most probable energy
b) mean energy
c) a maximum energy
a) most probable energy
b) mean energy
c) a maximum energy
Wednesday, March 5, 2008
1 mm Pb foil
1 mm Pb Foil is placed in the beam path during photon beam calibration because it:
introduces known electron contamination, facilitating the determination of PDD only due to photons, PDD (10)x, specifier of beam quality.
introduces known electron contamination, facilitating the determination of PDD only due to photons, PDD (10)x, specifier of beam quality.
Plane parallel chambers
Plane parallel chambers are recommended for the dosimetry of low-energy electron beams (<10 MeV) in place of Farmer type chambers, because:
Farmer type chambers give rise to significant electron fluence perturbation at low electron energies
Farmer type chambers are less accurrate for low-energy electrons.
Farmer type chambers give rise to significant electron fluence perturbation at low electron energies
Farmer type chambers are less accurrate for low-energy electrons.
Tuesday, March 4, 2008
TG51, Kq
Kq can vary from chamber to chamber by as much as 5%
http://www.irs.inms.nrc.ca/tg51/sld013.htm
http://www.irs.inms.nrc.ca/tg51/sld013.htm
Monday, March 3, 2008
AAPM TG-51 protocol
The AAPM TG-51 protocol is based on calibration in Co-60 beams in terms of
absorbed dose to water.
The reference phantom for beam calibration (AAPM TG-51) is water.
absorbed dose to water.
The reference phantom for beam calibration (AAPM TG-51) is water.
Recommended depth for photon beam calibration
The recommended depth for photon beam calibration is 10 cm.
The calibration depth must be much beyond the depth of dose maximum, ie. the transient equilibrium region, where the electron contamination does not reach.
A depth of 10 cm will satisfy this criterion for the whole range of clinically used photon beam qualities. Earlier recommendations gave depth of 5 cm, 7 cm and 10 cm for different beam qualities, but recent protocols recommend a single depth of 10 cm for all beam qualities.
The calibration depth must be much beyond the depth of dose maximum, ie. the transient equilibrium region, where the electron contamination does not reach.
A depth of 10 cm will satisfy this criterion for the whole range of clinically used photon beam qualities. Earlier recommendations gave depth of 5 cm, 7 cm and 10 cm for different beam qualities, but recent protocols recommend a single depth of 10 cm for all beam qualities.
Dmax for Co-60
The depth of dose maximum in a patient for a relatively clean Co-60 beam is
5 mm (0.5cm)
5 mm (0.5cm)
Surface or buildup region doses and measurement
Surface or buildup region doses must be measured using an extrapolation chamber.
Yes- The entry window thickness becomes the thickness of the dead layer of the skin and the chamber would measure the dose just below the dead layer of the skin.
Yes- The entry window thickness becomes the thickness of the dead layer of the skin and the chamber would measure the dose just below the dead layer of the skin.
Field size dependence of the PDD of an accelerator photon beam
The field size dependence of the PDD of an accelerator photon beam is very much influenced by the electron contamination in the beam.
This is because of electron and scatter photon contamination of the beam.
This is because of electron and scatter photon contamination of the beam.
AAPM TG-51 and electron beam quality
In AAPM TG-51 protocol, the electron beam quality is specified by the beam penetration parameter, R50,D which is the depth where the electron beam depth dose drops to 50% of the peak value. This is more accurate than the parameter, Eo, the mean energy of the electron beam used in the earlier protocols (as derived from R50 using an empirical equation)
Wednesday, February 27, 2008
Tuesday, February 26, 2008
Isocentric treatment technique
Isocentric treatment technique assures
Reproducibility of treatment setup.
Reproducibility of treatment setup.
Phantom scatter factor, Sp
Phantom scatter factor gives the
1) output factor measured in phantom
2) influence of head and phantom scatter on beam output measured in phantom with increasing collimator field size.
1) output factor measured in phantom
2) influence of head and phantom scatter on beam output measured in phantom with increasing collimator field size.
Collimator scatter factor
Collimator scatter Sc provideds the
1) output factor measured in air
2) influence of head scatter on beam output w ith increasing collimator field size
1) output factor measured in air
2) influence of head scatter on beam output w ith increasing collimator field size
TMR depends on
TMR depends on
depth
field size
beam quality
PDD depends on all of these plus distance to the surface (SSD). TMR was invented to simplify matters so they didn't have to reposition the patient as they rotate around as they did in SSD treatments with PDD.
depth
field size
beam quality
PDD depends on all of these plus distance to the surface (SSD). TMR was invented to simplify matters so they didn't have to reposition the patient as they rotate around as they did in SSD treatments with PDD.
TAR
TAR concept is usually not preferred with accelerator photon beams because the dose in free air is not well defined for these beams
TAR can be derived from PDD data
At dmax TAR and PSF are practically the same quantity.
TAR can be derived from PDD data
At dmax TAR and PSF are practically the same quantity.
Scatter and energy of the beam
At a phantom depth of 10 cm, on the central axis, scatter contribution to the dose is more for a Co-60 beam than an accelerator beam.
This is because scatter towards the central axis decreases as the beam quality increases; hence the decrease in the field size dependence for higher energy photon beams.
SUMMARIZED
At any given depth, the scatter contribution to the total dose decreases with an increase in incident photon energy.
This is because scatter towards the central axis decreases as the beam quality increases; hence the decrease in the field size dependence for higher energy photon beams.
SUMMARIZED
At any given depth, the scatter contribution to the total dose decreases with an increase in incident photon energy.
Off axis ratio
The off axis ratio is the dose at any point in a plane perpendicular to the central axis to that of the dose on the central axis in that plane.
Exit dose and the patient
Exit dose calculated using standard PDD tables will not give the correct exit dose. The actual dose will be less due to a lack of backscatter thickness at the patient exit point.
Output of a kV or a Co-60 unit
The output of a kV or a C0-60 unit can be measured in terms of air kerma rate.
TAR decreases with field asymmetry
TAR decreases with increased field asymmetry due to decrease in side scatter.
BSF and energy
At low energies (< about 0.5 mm Cu) BSF increases with an increase in energy
Due to increased penetration of back scattered radiation
At orthovoltage beam qualities (> about 0.5 mm Cu HVL) BSF actually decreases and forward scatter increases.
Due to increased penetration of back scattered radiation
At orthovoltage beam qualities (> about 0.5 mm Cu HVL) BSF actually decreases and forward scatter increases.
TAR
TAR can be measured directly or derived from PDD data.
TAR at dmax and PSF are the same quantity.
In general TAR increases with an increase in depth.
At shallow depths, TAR is greater for low beam qualities (kV x-ray region) because of increased side scatter.
At larger depths TAR is larger for higher energy photons compared to low beam qualities due to increased penetration.
TAR at dmax and PSF are the same quantity.
In general TAR increases with an increase in depth.
At shallow depths, TAR is greater for low beam qualities (kV x-ray region) because of increased side scatter.
At larger depths TAR is larger for higher energy photons compared to low beam qualities due to increased penetration.
TAR concept
TAR concept involves equilibrium does in air "free space dose" as defined in Khan 2003, which becomes more and more ambiguous and also difficult to evaluate.
For further study see Attix 1999. Note look for where the difficulty in the evaluation of this quantity for Co60 beams are discussed. So the TMR concept replaces this quantiy by a phantom defined quantity that is less ambiguous.
For further study see Attix 1999. Note look for where the difficulty in the evaluation of this quantity for Co60 beams are discussed. So the TMR concept replaces this quantiy by a phantom defined quantity that is less ambiguous.
Isodose and PDD data
Isodose and PDD data can't be used directly for a patient, without any modification. Often they have to be corrected for body curvature and internal heterogeneities since the beam data are acquired for a water phantom (homogeneous medium) for normal incidence.
Monday, February 25, 2008
Isocentric vs Constant SSD methods
In the case of isocentric treament, compared to constant SSD methods,
a) setup time decreases
b) setup error decreases
c) treatment accuracy increases
a) setup time decreases
b) setup error decreases
c) treatment accuracy increases
Disadvantage of using a physical wedge
Disadvantages of wedges oh there be a plenty:
1) hardening of the beam, which also depends on wedge thickness
2) attenuation of the beam and an increase in treatment tme
3) inconvenience in handling heavy wedges. Gropes and complaining from therapists who haven't seen the inside of a gym since about the time Jane Fonda released her first fitness video on VHS tape.
1) hardening of the beam, which also depends on wedge thickness
2) attenuation of the beam and an increase in treatment tme
3) inconvenience in handling heavy wedges. Gropes and complaining from therapists who haven't seen the inside of a gym since about the time Jane Fonda released her first fitness video on VHS tape.
Independent movement of jaws
Independent movement of jaws is necessary for
a) creating of blocked fields
b) avoiding beam divergence at one edge when necessary
c) creating dynamic wedge fields.
a) creating of blocked fields
b) avoiding beam divergence at one edge when necessary
c) creating dynamic wedge fields.
A beam flattening filter is used with
A beam flattening filter is used with a
PHOTON BEAM
Now who is smart than a 5th grader. Most 5th grade medical physicists would know this answer.
Followup
How is it retracted for e-beams?
I know it's on a carousel.
PHOTON BEAM
Now who is smart than a 5th grader. Most 5th grade medical physicists would know this answer.
Followup
How is it retracted for e-beams?
I know it's on a carousel.
Low energy electrons and scattering
Low energy electrons are more easily scattered compared to high energy electrons.
Yes remember Scattering is Zsquared/Esquared
Yes remember Scattering is Zsquared/Esquared
SCIENTIFIC NOTATION!!!!!
For $%$%$sake, why doesn't blogger allow scientific notation. Has our society devolved to where all the blogs revolve around mere words, the tabloids, gossip and political bullshit?
What happened to people who want to use this for their mathematical equations? Am i the only one out there.
Buehler, Buehler.....
What happened to people who want to use this for their mathematical equations? Am i the only one out there.
Buehler, Buehler.....
Electron beams and their range
Electron beams have a finite range in the medium.
The range of electons of energy E (MeV) in water is roughly E/2 cm.
NOTE: The range of electrons in lead is E/20 cm (per WMK and Khan e- chapter)
The range of electons of energy E (MeV) in water is roughly E/2 cm.
NOTE: The range of electrons in lead is E/20 cm (per WMK and Khan e- chapter)
energy fluence of an accelerator electron beam.
Energy fluence of an accelerator electron beam and dose fall off.
IT DOSE NOT FALL OFF ACCORDING THE ISL. The dose measured at dmax point in a phantom, by varying the SSD, roughly follows (to better than 2%) the ISL for small changes in distance (say within about 5 cm), but not with respect ot the target position, but with respect to a point much closer than the target. This apparent distance, which is also a function of field size, must be experimentally determined for every electron beam quality of the machine (See Khan 2003 for more details)
IT DOSE NOT FALL OFF ACCORDING THE ISL. The dose measured at dmax point in a phantom, by varying the SSD, roughly follows (to better than 2%) the ISL for small changes in distance (say within about 5 cm), but not with respect ot the target position, but with respect to a point much closer than the target. This apparent distance, which is also a function of field size, must be experimentally determined for every electron beam quality of the machine (See Khan 2003 for more details)
Sunday, February 24, 2008
Megavoltage x-ray vs kV
kV imaging utilizes a different method of filtration to attenuate the beam. The filter may be something like aluminum, ex 200 kVp x-ray beam filtered by 1 mm thick aluminum filter. This distribution includes the glass envelope of the x-ray tube, the surrounding oil, and the exit window of the tube housing as well. This so called inherent filtration is equivalent to approximately 1 mm Al in most x-ray tubes. Can also be a combination filter such as a Thoraeus filter TCA (tin copper aluminum)
For a megavoltage x-ray beam however, the beam is hardened by the inherent filtration of the transmission target as well as by transmission through the flattening filter.
For a megavoltage x-ray beam however, the beam is hardened by the inherent filtration of the transmission target as well as by transmission through the flattening filter.
Buildup in electron beams
Buildup in electron beams is due to electron scattering and the finite range of the delta rays produced.
BONUS NOTE
The tail of the electron beam depth dose distribution is due to dose produced by the brehmmstrahlung radiation or the x-ray background (photon contamination)
BONUS NOTE
The tail of the electron beam depth dose distribution is due to dose produced by the brehmmstrahlung radiation or the x-ray background (photon contamination)
Scatter and PDD
Scatter generally increases PDD. Since it generally increases dose on the central axis.
Beam divergence and beam attenuation
Beam divergence and beam attenuation in a phantom decrease with PDD.
Yese because both parameters reduce the dose at a given depth.
Yese because both parameters reduce the dose at a given depth.
Surface buildup and extrapolation chamber
The entry window thickness becomes the thickness of the dead layer of the skin and the chamber would measure the dose just below the dead layer of the skin.
Field size dependence and electron contamination
The field size dependence of the PDD of an accelerator photon beam is very much influenced by the electron contamination of the beam.
This is becausee of electron and scatter photon contamination of the beam.
This is becausee of electron and scatter photon contamination of the beam.
AAPM TG-51 protocol
In the AAPM TG-51 protocol for electrons, Electron beam quality is specified by the beam penetration parameter R50,d, the depth where the electron beam depth dose drops to 50% of the peak value. This is more accurate than the parameter Eo, the mean energy of the electron beam used in the earlier protocols (as derived from R50 using an empirical equation).
AAPM TG-51 and photons
In the AAPM TG-51 protocol, beam quality for photons is specified by PDD (10)x.
Photons and Inverse Square Law (ISL)
The energy fluence of the photons fall off according to the Inverse square law of distance with the source or the target position as the origin.
The dose measured at the dmax point in a phantom by varying the SSD, roughly follows (to better than 2%) the inverse square law, for moderate changes in SSD (say within about 20 cm). This however must be verified for the treatment machine before putting it into clinical use.
SIDE NOTE:
Note that photons do not have a finite range in the patient.
The dose measured at the dmax point in a phantom by varying the SSD, roughly follows (to better than 2%) the inverse square law, for moderate changes in SSD (say within about 20 cm). This however must be verified for the treatment machine before putting it into clinical use.
SIDE NOTE:
Note that photons do not have a finite range in the patient.
Dynamic Wedge
A dynamic wedge is created by moving one jaw of the collimator towards the other, which creates the wedged field profile.
One of the advantages of a dynamic wedge over a physical wedge for treatment is that there is no change in beam quality across the field size. (No beam hardening for a dynamic wedge like there is when a physical wedge differentially attenuates the photon beam).
One of the advantages of a dynamic wedge over a physical wedge for treatment is that there is no change in beam quality across the field size. (No beam hardening for a dynamic wedge like there is when a physical wedge differentially attenuates the photon beam).
UNIVERSAL WEDGE
A universal wedge is an internally mounted wedge that is usable for many different field sizes.
Scatter contribution
At any given depth, the scatter contribution to total dose decreases with increase in photon energy.
CORRECT (SEE EARLIER POST)
CORRECT (SEE EARLIER POST)
Flattening filter
C0-60 machines do not have a flattening filter because the photon emission from a Co-60 source is more isotropic compared to high energy photon beams, which are peaked in the direction of a central axis. So a flattening filger is not necessary.
Scatter contribution to dose
At a phantom depth of 10 cm, on the central axis, scatter contribution to the dose is more for an accelerator photon beam compared to a Co-60 beam.
NO, Scatter towards the central axis decreases as the beam quality (energy) increases; hence the decrease in the field size dependence seen for higher energy photon beams.
SO SCATTER CONTRIBUTION TO DOSE IS LESS FOR LINACS THAN A CO-60 Machine.
NO, Scatter towards the central axis decreases as the beam quality (energy) increases; hence the decrease in the field size dependence seen for higher energy photon beams.
SO SCATTER CONTRIBUTION TO DOSE IS LESS FOR LINACS THAN A CO-60 Machine.
Off-axis ratio
The off-axis ratio (OAR) is the dose at any point in a plane perpendicular to the central axis to that of the dose on the central axis in that plane.
Off-axis ratios can be used to determine dose at off axis points.
Off-axis ratios can be used to determine dose at off axis points.
Output of kV or Co-60 units
The output of a KV or a Co-6o unit can be measured in terms of air kerma rate.
BSF (Back Scatter Factor)
At low energies (< about 0.5 mm Cu) BSF increases with an increase in energy.
Yes, this occurs due to increased penetration of backscattered radiation.
At orthovoltage beam qualities (> about 0.5 mm Cu HVL)
BSF actually decreases and forward scatter increases.
Yes, this occurs due to increased penetration of backscattered radiation.
At orthovoltage beam qualities (> about 0.5 mm Cu HVL)
BSF actually decreases and forward scatter increases.
TAR (tissue air ratio)
TAR concept is usually not used with the accelerator photon beams because the dose in free air is not well defined due to conceptual difficulties.
TAR concept involves equilibrium dose in air ("free space dose") as defined in Khan (2003), which becomes more and more ambiguous and also difficult to evaluate.
For further study (see Attix 1999) Note look for where the difficulty in the evaluation of this quantity for Co-60 beams is discussed. So the TMR concept replaces this quantity by a phantom defined quantity that is less ambiguous.
TAR can be measured directly or derived from the PDD data.
TAR at dmax and PSF (peak scatter factor) are the same quantity.
In general TAR decreases with an increase in depth.
At shallow depths TAR is greater for low beam qualities (kV x-ray region) because of increased side scatter.
At larger depths TAR is larger for high energy photons compared to low beam energies, due to increased penetration.
TAR increases with an increase in field size.
TAR decreases with increasing field asymmetry. This occurrs due to decrease in side scatter.
TAR concept involves equilibrium dose in air ("free space dose") as defined in Khan (2003), which becomes more and more ambiguous and also difficult to evaluate.
For further study (see Attix 1999) Note look for where the difficulty in the evaluation of this quantity for Co-60 beams is discussed. So the TMR concept replaces this quantity by a phantom defined quantity that is less ambiguous.
TAR can be measured directly or derived from the PDD data.
TAR at dmax and PSF (peak scatter factor) are the same quantity.
In general TAR decreases with an increase in depth.
At shallow depths TAR is greater for low beam qualities (kV x-ray region) because of increased side scatter.
At larger depths TAR is larger for high energy photons compared to low beam energies, due to increased penetration.
TAR increases with an increase in field size.
TAR decreases with increasing field asymmetry. This occurrs due to decrease in side scatter.
Isodose distribution and proper phantom size
To measure the isodose distribution for a 10cmx10cm field size, it is not necessary to use a 40x40 water phantom.
The requirement is that the field must be surrounded by 5 cm of water on all sizes. So for smaller field sizes, smaller phantoms can be made use of.
The requirement is that the field must be surrounded by 5 cm of water on all sizes. So for smaller field sizes, smaller phantoms can be made use of.
SSD vs SAD techniques
In SSD techniques, the beams are weighted at dmax points
In SAD techniques, teh beams are weighted at the target center.
In SAD techniques, teh beams are weighted at the target center.
Exposure of 1 roentgen
An exposure of 1 roentgen corresponds to a charge release of
1 esu
also 2.58x10 -4 C/kg
1 esu
also 2.58x10 -4 C/kg
Farmer chamber requirement
A Farmer type chamber used for exposure or air kerma measurements in a hospital must have a tissue equivalent wall.
Farmer chambers and buildup cap
Farmer chambers are usually calibrated with a buildup cap at Co-60 energy to
PROVIDE EQUILIBRIUM WALL THICKNESS
PROVIDE EQUILIBRIUM WALL THICKNESS
Equilibrium wall thickness of chambers
KV region and chambers
The wall thickness of the chamber used for beam output in the kV range is about 60 to 90 mg/cc
The equilibrium wall thickness required for the chamber used for beam output measurements at Co-60 energy is about 500 mg/cc.
The wall thickness of the chamber used for beam output in the kV range is about 60 to 90 mg/cc
The equilibrium wall thickness required for the chamber used for beam output measurements at Co-60 energy is about 500 mg/cc.
Plane parallel chambers (and perturbation)
Some plane parallel chambers exhibit significant perturbation at low electron energies.
TRUE.
For example, the Markus chamber does.
The Roos chamber designed by PTW Germany is a modification over the Markus Chamber and does not exhibit fluence perturbation at low electron energies. It's stability however depends on the reliability of the aquaduc coating on it.
TRUE.
For example, the Markus chamber does.
The Roos chamber designed by PTW Germany is a modification over the Markus Chamber and does not exhibit fluence perturbation at low electron energies. It's stability however depends on the reliability of the aquaduc coating on it.
Farmer type chambers
Farmer type chambers can be used for the dosimetry of accelerator produced photon and electron beams.
TRUE (except for incident electron beam energies with Eo<10 MeV).
So why do we use these for e- measurements? Relative is OK but what about absolute?
TRUE (except for incident electron beam energies with Eo<10 MeV).
So why do we use these for e- measurements? Relative is OK but what about absolute?
Chamber calibration
Hospital chambers must be calibrated in terms of exposure or air kerma or absorbed dose to water in an Accredited Dosimetry Calibration Laboratory, which is traceable to the Primary Standards Laboratory.
Graphite cavity ionization chambers
Graphite cavity ionization chambers are the primary standard of absorbed dose for therapy machines.
For kv x-ray beams, free air ionization chambers are the primary standard of exposure or air kerma.
For kv x-ray beams, free air ionization chambers are the primary standard of exposure or air kerma.
Primary Standards of Exposure
Primary Standards of Exposure, Air Kerma or Water Absorbed Dose have been established by some Primary Standards Dosimetry Laboratories (PSDL's) for a Co-60 beam.
Hospital chambers used for beam calibration must be traceable to these standards either directly or indirectly.
Hospital chambers used for beam calibration must be traceable to these standards either directly or indirectly.
Friday, February 22, 2008
Ionization curves to depth dose curves.
How do you convert ionization curves to % depth dose curves?
Ref p 303 Khan
Depth ionization curves obtained with air ionization chambers can be converted into depth dose curves by making corrections for change in stopping power ratio of water to air with depth.
Ref p 303 Khan
Depth ionization curves obtained with air ionization chambers can be converted into depth dose curves by making corrections for change in stopping power ratio of water to air with depth.
TG-51 Effective point of measurement of parallel plate chamber.
Where is the effective point of measurement of parallel plate chambers?
For plane parallel chambers, the center of the front (upstream) face of the chamber air cavity is the point of measurement.
For plane parallel chambers, the center of the front (upstream) face of the chamber air cavity is the point of measurement.
Purpose of the guard ring in the plane parallel chamber is to
The purpose of the guard ring in the plane parallel chamber is to:
(DEFINE THE COLLECTION VOLUME SEEMED TO BE THE ONLY REASONABLE ANSWER, CHECK THIS.)
A guard ring prevents leakage. It creates a homogeneous electric field between the two electrodes. It prevents secondary electrons scattered from the wall being counted in the chamber.
Add picture from Pgorsak.
(DEFINE THE COLLECTION VOLUME SEEMED TO BE THE ONLY REASONABLE ANSWER, CHECK THIS.)
A guard ring prevents leakage. It creates a homogeneous electric field between the two electrodes. It prevents secondary electrons scattered from the wall being counted in the chamber.
Add picture from Pgorsak.
TG-51 Cross calibrating parallel plate chamber
To cross calibrate a parallel plate chamber what would one use?
OPTIONS: Co60, high energy photons, high energy electrons, low energy electrons.
Per tG-51 pg 1860.
Since Co-60 cal factors of plane parallel chambers are sensitive to their construction,
calibrate against calibrated cylindrical chambers in a high energy electron beam.
OPTIONS: Co60, high energy photons, high energy electrons, low energy electrons.
Per tG-51 pg 1860.
Since Co-60 cal factors of plane parallel chambers are sensitive to their construction,
calibrate against calibrated cylindrical chambers in a high energy electron beam.
How many TVL's in a linac head
At 1 m from the source, nonshielded vs shielded you get 0.1% (leakage through the head).
To get that think of half values 1/2=50 .......1/16=6.25.........1/1024=10 HVL'S
So 10 HVL's are required to get it to 0.1%
We know 1 TVL=3.3 HVL's so, so to get 10 HVL's we need about 3 TVL's.
The answer is then 3 TVL's.
To get that think of half values 1/2=50 .......1/16=6.25.........1/1024=10 HVL'S
So 10 HVL's are required to get it to 0.1%
We know 1 TVL=3.3 HVL's so, so to get 10 HVL's we need about 3 TVL's.
The answer is then 3 TVL's.
Tuesday, February 19, 2008
Regulations
Received a link to this website today. It's quite useful for linking you to various regulations
http://www.physics.isu.edu/radinf/rsotoolbox.htm
http://www.physics.isu.edu/radinf/rsotoolbox.htm
Absorbed Dose to Water
Therapy chambers can be calibrated in terms of exposure, air kerma or "Absorbed Dose to Water". Accredited laboratories offer such calibrations.
Farmer ionization chamber
A Farmer type ioniziation chamber must be used for radiation therapy beam calibration.
Volume about 0.6 cc cubed.
Volume about 0.6 cc cubed.
Exposure is the ionization equivalent of collision kerma in air
Exposure is the ionization equivalent of collision kerma in air. In the figure above (ADD), exposure at
point P= X(P)= Kc(P)/We) (J/kg) (J/c)
To obtain exposure in roentgens we must use the conversion factor 2.58 x10 -4 (C/kg)/R
point P= X(P)= Kc(P)/We) (J/kg) (J/c)
To obtain exposure in roentgens we must use the conversion factor 2.58 x10 -4 (C/kg)/R
Collision kerma Kc
That part of kerma (i.e. electron kinetic energy) that leads to subsequent collision interactions and hence energy deposition is known as collision kerma (Kc).
Collision part of kerma is NOT ALWAYS equal to absorbed dose. But when charged particle equilbrium exists, for example, at P in the above figure, electron energy escaping from at a point P will be compensated by electron energy entering at P (ENTER DIAGRAM LATER). In this case Kc(P) can be equated to D(P) with a small correction that accounts for the very small attenuation of photons over electron ranges.
Since any dosimetric system, whether it is a Primary Standard or a therapy dosimetry responds to energy absorbed in the sensitive volume, kerma (or exposure) of indirectly ionizing particles can be measured only under electronic equilibrium conditions.
Collision part of kerma is NOT ALWAYS equal to absorbed dose. But when charged particle equilbrium exists, for example, at P in the above figure, electron energy escaping from at a point P will be compensated by electron energy entering at P (ENTER DIAGRAM LATER). In this case Kc(P) can be equated to D(P) with a small correction that accounts for the very small attenuation of photons over electron ranges.
Since any dosimetric system, whether it is a Primary Standard or a therapy dosimetry responds to energy absorbed in the sensitive volume, kerma (or exposure) of indirectly ionizing particles can be measured only under electronic equilibrium conditions.
Kerma and energy deposition
Kerma is only defined for photons which interact with matter and convert to electron energy.
Not all of kerma ends up as energy deposition in the medium. Part of kerma (electron kinetic energy produced) is converted into bremmstrahlung which escapes from the region of interest, so that only that part of kerma called the collision part of kerma is energy that is dissipated in the medium by ionization and excitation events.
Not all of kerma ends up as energy deposition in the medium. Part of kerma (electron kinetic energy produced) is converted into bremmstrahlung which escapes from the region of interest, so that only that part of kerma called the collision part of kerma is energy that is dissipated in the medium by ionization and excitation events.
Kerma
Kerma, K, at a particular point is defined as the kinetic energy released (ex in the form of charged particle energies) per unitmass, in an infinitesmal volume, around that point by IDIR (indirectly ionizing radiation)
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