12 Physics and Clinical Aspects of Brachytherapy Z

12 Physics and Clinical Aspects of Brachytherapy

Zuofeng Li

Z. Li, DSc

Department of Radiation Oncology, Washington University, St. Louis, MO 63110, USA

12.1

Introduction

Brachytherapy is the use of sealed radioactive sources placed in close proximity to the treatment target volume, either by directly inserting them into the tumor, or by loading them into instru- ments (applicators) which were previously inserted into cavities inside the body at close distance to the tumor. Brachytherapy may be used as a sole radiation treatment of the tumor, such as in the case of early-stage prostate and breast cancers. It is also often used in combination with external beam radiation therapy to deliver a boost radiation dose to the tumor, as in the case of gynecological tumors, later-stage prostate cancer, and many head and neck cancers. Following the surgical removal of the gross tumor, brachytherapy may be used to

CONTENTS

12.1 Introduction 255

12.2 Classifications of Brachytherapy 256 12.2.1 Permanent Versus Temporary Implants 256 12.2.2 Interstitial, Intracavitary/Intraluminal, Topical/Mold 256

12.2.3 Hot Loading, Manual Afterloading, and Remote Afterloading 256

12.2.4 Low Dose Rate, High Dose Rate,

Medium Dose Rate, and Pulse Dose Rate 257 12.3 Physical Characteristics of

Brachytherapy Sources 258 12.3.1 Half-Life 258

12.3.2 Specific Activity 258 12.3.3 Average Energy 259

12.4 Sources Used in Brachytherapy 259 12.4.1 High-Energy Photon Emitters 259 12.4.2 Low-Energy Photon Emitters 260 12.4.3 Emerging Sources 261

12.5 Dose Calculations in Brachytherapy 261 12.5.1 The Superposition Principle 261 12.5.2 Source Strength Units 261 12.5.3 Single Source Dosimetry 262 12.5.3.1 Point Source Dosimetry 262 12.5.3.2 Line Source Dosimetry 263

12.5.3.3 Total Delivered Dose Calculations 265 12.6 Gynecological Intracavitary Implant 266 12.6.1 Applicators 266

12.6.2 Dose Specification for

Cervical Cancer Treatments 266 12.6.3 ICRU Report 38 Recommendations 266

12.6.4 Volumetric Image-Based GYN Brachytherapy 266 12.7 Interstitial Implant Dosimetry Systems 268 12.7.1 Paterson–Parker (Manchester) System 269 12.7.2 Quimby System 269

12.7.3 Paris System 269

12.7.4 ICRU Report 58 Recommendations 270

12.8 Process of Brachytherapy Treatment Planning and Delivery 270

12.8.1 Preplanning 270

12.8.2 Source and Applicator Preparation 271 12.8.3 Applicator and Catheter Insertion 271 12.8.4 Source/Applicator Localization 272 12.8.5 Treatment Planning and Quality Assurance

Review of Brachytherapy Treatment Plan 273

12.8.6 Source Loading and Treatment Delivery 274 12.9 Transrectal Ultrasound-Guided Permanent Prostate Brachytherapy 274

12.9.1 Dosimetric Goals of a Permanent Prostate Brachytherapy Treatment 274 12.9.2 Equipment 275

12.9.3 Volume Study 275 12.9.4 Treatment Planning 276 12.9.5 Treatment Plan Review and

Pretreatment Quality Assurance Tests 277 12.9.6 Implant Procedure 278

12.9.7 Post-Implant Dosimetry 278 12.10 High Dose Rate Remote-Afterloading Brachytherapy 279

12.10.1 Equipment and Operating Principles 280 12.10.2 Clinical Application of HDR Brachytherapy:

Interstitial Accelerated Partial Breast Irradiation 281 12.11 Radiation Protection and Regulatory Compliance in Brachytherapy 285

12.11.1 Licensing, Authorization, and Report of Medical Events 286

12.11.2 Radiation Safety Concerns in Permanent Implants 287 12.11.3 Safety and Regulatory Issues in HDR Brachytherapy 287 References 288

delivery a tumoricidal radiation dose to the tumor bed, where microscopic diseases remain. Due to the rapid falloff of dose away from the sources, brachy- therapy allows the delivery of greater tumor doses than external beam radiation therapy, while retain- ing excellent sparing of neighboring critical organs.

Compared with surgery, brachytherapy does not create a tissue deficit, thereby allowing potentially better cosmetic results.

12.2

Classifications of Brachytherapy

Brachytherapy modalities can be classified according to various criteria, including implant duration, the approach used to insert the sources into the patient, the technique used to load the sources, and the rate at which radiation dose is delivered to the target. These classifications hold significance not only as medical terms, but also in the selection of radioactive sources for a given brachytherapy treatment.

12.2.1

Permanent Versus Temporary Implants

Permanent brachytherapy implants are those where the sources are inserted into the patient, remain- ing permanently in the patient. Common permanent brachytherapy procedures include treatment of the prostate, head and neck cancers, lung, and sarco- mas. I-125 and Pd-103 seeds are commonly used for permanent implants (Nag et al. 1999, 2000), although Au-198 seeds have also been used occa- sionally (Crusinberry et al. 1985; Hochstetler et al. 1995). As is discussed later, sources used for per- manent implants need to have low energy, short half- lives, or a combination of both, so that the radiation exposure received by people that have either frequent or close contact with the patient is limited.

Temporary brachytherapy implants are those where the sources are implanted in the patient for a pre-determined length of time and then removed.

Treatment times of temporary implants range from a few minutes, when the high dose rate afterload- ing technique is used, to a few days, for common low dose rate treatments. Patients may need to be admitted into the hospital for the duration of the treatment. Radiation exposure to hospital workers is therefore a significant concern when temporary implants are employed.

12.2.2

Interstitial, Intracavitary/Intraluminal, Topical/Mold

Depending on the approach used to insert the brachytherapy sources into the patient, brachyther- apy can be classified into interstitial, intracavitary/

intraluminal or topical/mold treatments.

In interstitial brachytherapy treatments, brachy- therapy sources are introduced into the tissue, often with the use of needles and catheters of small diam- eters, used to minimize trauma to the normal tissue.

Correspondingly, brachytherapy sources used for interstitial treatment need to have small dimen- sions to fit into the needles and catheters. Intersti- tial treatments are used for tumors such as prostate cancer, breast cancer, and sarcomas.

In intracavitary brachytherapy treatments, sources are loaded into applicators, which are posi- tioned into cavities within the human anatomy adjacent to the target tissue. Treatment site-spe- cific applicators are designed to fit into cavities and place the sources near the target tissues. Examples include the tandem and ovoid applicators for treat- ment of cervical cancer, the cylinder applicator for treatment of vaginal cancer, and the nasopharyngeal applicator for treatment of cancer of the nasophar- ynx. The applicators remain in the patient during the treatment and are removed at the completion of the treatment, so intracavitary brachytherapy treat- ments are usually temporary treatments.

12.2.3

Hot Loading, Manual Afterloading, and Remote Afterloading

Depending on the timing of source insertion rela-

tive to the surgical procedure to insert the applica-

tors and/or needles, brachytherapy can be divided

into hot loading, in which the sources are inserted

in the operation room (OR) immediately after the

applicators are inserted; manual afterloading, where

the applicators are inserted into the patient in the

OR, and sources are loaded after the patient’s return

from the recovery room to the patient’s room; and

remote afterloading, in which a computer-con-

trolled device is used to load the sources automati-

cally, thus eliminating manual handling of radioac-

tive sources altogether. Hot loading is rarely used

currently due to the high radiation exposure to OR,

recovery room, and transportation personnel when

compared against afterloading. Permanent implant

treatments are typically hot-loaded, as are pre-fab- ricated eye plaque applicators containing

I or

103

Pd seeds for treatment of ocular melanoma. Use of afterloading techniques, either manual or remote (computer controlled), minimizes radiation exposure to hospital personnel, in addition to providing an opportunity for the treatment planner to optimize the source strength and loading distribution based on a retrospective review of the applicator position- ing relative to the target tissue.

12.2.4

Low Dose Rate, High Dose Rate, Medium Dose Rate, and Pulse Dose Rate

Brachytherapy treatments can also be classified according to the dose rate at which brachytherapy treatments are delivered:

1. Low dose rate (LDR): D

< 120 cGy/h

2. Medium dose rate (MDR):

120 cGy/h d

<1200 cGy/h 3. High dose rate (HDR): D

≥ 1200 cGy/h

Much of the existing clinical brachytherapy expe- rience was for treatments delivered using the classic LDR regimen, at D

≈45 cGy/h. The biological effec- tiveness of brachytherapy treatment depends sig- nificantly on the dose rate at which the treatment is delivered. Much effort has been spent, therefore, on the biological effect of brachytherapy delivery at higher dose rates, such that a dose biologically equivalent to previous LDR treatments can be deliv- ered. Remote afterloading HDR units, equipped with a high-activity

Ir source, can deliver an entire treatment fraction in minutes. At such high dose rates, the advantage of normal tissue repair associated with LDR is lost, so HDR treatments must be fractionated, delivering a smaller total dose rela- tive to their LDR counterparts. This has led to the

Table 12.1 Common brachytherapy sources and their physical characteristics. (From Williamson 1998b). LDR low dose rate, HDR high dose rate

Element Isotope Energy (MeV)

Half-life HVL- lead (mm)

Exposure rate con- stant (*_G)^a

Source form Clinical application

Obsolete sealed sources of historic significance

Radium ²²⁶Ra 0.83 (avg) 1626 years 16 8.25^b Tubes and needles LDR intracavitary and interstitial Radon ²²²Rn 0.83 (avg) 3.83 days 16 8.25^b Gas encapsulated in

gold tubing

Permanent interstitial; temporary molds

Currently used sealed sources

Cesium ¹³⁷Cs 0.662 30 years 6.5 3.28 Tubes and needles LDR intracavitary and interstitial Iridium ¹⁹²Ir 0.397 (avg) 73.8 days 6 4.69 Seeds in nylon ribbon;

metal wires; encapsu- lated source on cable

LDR temporary interstitial; HDR interstitial and intracavitary Cobalt ⁶⁰Co 1.25 5.25 years 11 13.07 Encapsulated spheres HDR intracavitary

Iodine ¹²⁵I 0.028 59.6 days 0.025 1.45 Seeds Permanent interstitial

Palladium ¹⁰³Pd 0.020 17 days 0.013 1.48 Seeds Permanent interstitial

Gold ¹⁹⁸Au 0.412 2.7 days 6 2.35 Seeds Permanent interstitial

Strontium ⁹⁰Sr–⁹⁰Y 2.24 βmax 28.9 years – – Plaque Treatment of superficial ocular lesions

Developmental sealed sources

Americium ²⁴¹Am 0.060 432 years 0.12 0.12 Tubes LDR intracavitary

Ytterbium ¹⁶⁹Yb 0.093 32 days 0.48 1.80 Seeds LDR temporary interstitial Californium ²⁵²Cf 2.4 (avg)

neutron

2.65 years – – Tubes High-LET LDR intracavitary

Cesium ¹³¹Cs 0.030 9.69 days 0.030 0.64 Seeds LDR Permanent implants

Samarium ¹⁴⁵Sm 0.043 340 days 0.060 0.885 Seeds LDR temporary interstitial Unsealed radioisotopes used for radiopharmaceutical therapy

Strontium ⁸⁹Sr 1.4 βmax 51 days – – SrCl², IV solution Diffuse bone metastases Iodine ¹³¹I 0.61 βmax

0.364 MeV γ

8.06 days – – Capsule NaI oral solu- tion

Thyroid cancer Phosphorus ³²P 1.71 βmax 14.3 days – – Chromic phosphate

colloid instillation;

Na₂PO₂ solution

Ovarian cancer seeding; peri- toneal surface; PVC, chronic leukemia

a No filtration in units of R × cm² × mCi^–1 × hr^{–1 b} 0.5-mm platinum filtration; units of R/cm²/mCi^–1/hr

development of the pulsed dose rate delivery, in which the overall treatment time is equivalent to a traditional low dose rate treatment at 40–80 h. The sources, however, are only inserted into the patient for minutes during each hour of treatment, resulting in higher instantaneous dose rate, through the same dose delivered within each hour as a traditional LDR treatments. Several authors (Brenner et al. 1996, 1997; Chen et al. 1997; Visser et al. 1996) have dem- onstrated biological equivalence of PDR relative to LDR treatments.

12.3

Physical Characteristics of Brachytherapy Sources

A brachytherapy source is characterized by the rate at which its strength decays (half-life), by how much radioactivity can be obtained for a given mass of the radioactive source (specific activity), and by the energies and types of the radiation particles that are emitted from the source (energy spectrum).

These physical brachytherapy source characteristics will guide the clinical utilization. Table 12.1 lists the common radioactive sources used in brachytherapy, together with their physical characteristics.

12.3.1 Half-Life

The strength of a radiation source decays exponen- tially. Let the strength of the source at time 0 be A

. The strength of the source A(t) at time t is then given by the equation

A(t) = A₀

× e

, (1)

where µ is the decay constant. µ describes the rate at which the source strength decays. Of particular use in brachytherapy is the time it takes for the source strength to decay to half of its initial value, i.e., A(T

) = A

/2, where T

is the half-life of the source. Substituting into Eq. (1), we obtain

A(T_1/2

) = A

/2 = A

× e

; (2) or

. (2a)

A source’s half-life is a fundamental quantity of the radioactive nuclide of the source. Common brachytherapy sources have half-lives ranging from days to years. The length of a given brachytherapy source’s half-life determines its shelf life, namely, whether a source can be stored and used repeatedly over a long period of time. Sources with shorter half- lives, such as

I and

Pd sources, need to be pur- chased and received with an accurate knowledge of the source strength relative to the intended implant procedure date, so that the source strength on the day of implant is as prescribed, and that the desired initial dose rate (in terms of cGy/h or cGy/day) is achieved. Sources with longer half-lives, such as

137

Cs and

Ir sources, can be used for treatment of multiple patients before replacement, thereby reduc- ing the cost of each treatment.

A source’s half-life, together with its average energy, determines its suitability for use in perma- nent or temporary implants. When brachytherapy sources are permanently inserted into a patient and the patient is released from the hospital, the radia- tion exposure around the patient can pose a risk to people that are present within short distances from the patient. Sources with shorter half-lives can reduce these risks because the radiation exposure around the patient decreases rapidly with time. If necessary, the patient can be hospitalized in a pri- vate room for a short period of time.

The half-life of a brachytherapy source also impacts the implant dose calculation. The decay of the source may not need to be explicitly accounted for if the source has a suffi ciently long halfl ife. For example,

Cs sources, with a half-life of 30 years, may be assumed to hold a constant source strength during the treatment period of a few days, whereas the dose calculation of an implant using

I sources, with a half-life of 59.8 days, needs to consider the decay of the sources during the implant.

12.3.2

Specific Activity

The strength of a brachytherapy source for practical

applications is limited by its specific activity. The spe-

cific activity is the ratio of activity contained within

a unit mass of the source. When a parent nuclide is

activated within a neutron flux field, the number

of radioactive nuclides per unit mass that may be

obtained is limited by the neutron flux field strength,

the parent nuclide’s neutron cross section, and the

source half-life. This is important for HDR intersti-

tial brachytherapy applications, which require small source dimensions as well as high source strengths.

The popularity of the

Ir source in modern brachy- therapy is partly due to its high specific activity and high neutron cross section, thereby making it suitable as an HDR remote afterloading source. The small size of the source makes it useful for both interstitial and intracavitary brachytherapy treatments.

12.3.3

Average Energy

The average energy of a brachytherapy source deter- mines the penetrability of the photon particles emitted from the source. The high-energy photon sources allow higher radiation dose to tissues at larger distances to the sources, such as the pelvis nodes in the treatment of cervical cancer. On the other hand, the high-energy photons require thicker shields for protection of hospital personnel. Perma- nent brachytherapy treatments often use low-energy photon emitting sources, such as

I and

Pd, as the photons from these sources are mostly atten- uated by the patient tissue, resulting in very low radiation exposure rates around the patient. Patients treated with these sources can be released from the hospital without violating federal regulations on radiation exposure to members of the public from the implanted sources. When high-energy sources, such as

Au, are used for permanent implants, the patient needs to be confined in the hospital until the source strength decays to a suitable value, such that the radiation exposure from the sources out- side the patient satisfies the limits of these regula- tions. For these reasons,

Rn and

Au are the only sources useful for permanent brachytherapy implant because their short half-lives of approximately 3 days allow adequate source decay during the patient’s hospital stay.

I and

Pd sources, however, can be easily shielded by a thin lead foil, making them useful for treatments of shallowly located or super- ficial tumors such as ocular melanoma.

12.4

Sources Used in Brachytherapy

The use of radioactive sources for treatment of malig- nancies started shortly after the discovery of radium in 1898 by Madame Curie.

Ra, sealed in platinum tubes or needles, was used for interstitial and intra-

cavitary temporary treatments.

Rn, the daughter product of

Ra, in a gas form sealed within a gold seed, was later used for permanent implants, due to its short half-life. While neither sources are currently used clinically, much of the current brachytherapy treatments derive the dose specification and pre- scription parameters from the earlier clinical experi- ences using

Ra and

Rn sources. Their historical importance therefore cannot be ignored.

12.4.1

High-Energy Photon Emitters

The following information is given about high- energy photon emitters:

1. Radium-226:

Ra was the fi rst radionuclide iso- lated, and the fi rst used in brachytherapy treat- ments. Radium-226 has a half-life of 1620 years.

The J-rays from radium and its decay products range in energy from 0.05 to 2.4 MeV, with an average energy of about 0.8 MeV. The active

Ra sources consist of a radium salt (sulfate) mixed with fi ller (usually barium sulfate), which is encap- sulated in platinum cylinders to form radium tubes or needles. Radium tubes have a platinum wall thickness of 0.5 mm, are typically 22 mm long, and contain from 5 to 25 mg of radium in 15-mm active lengths. Radium tubes have a wall thickness of 1.0 mm and are often classifi ed their strengths per centimeters of length. Full-intensity needles typically have 0.66 mg of radium per centime- ter of length; half-intensity needles have 0.33 mg of radium per centimeter, and quarter-intensity needles have 0.165 mg of radium per centimeter.

2. Radon-222:

Rn, with a half-life of 3.83 days and average energy of 1.2 MeV, is a gas produced when radium decays. The radon gas was extracted and encapsulated in gold seeds, which were used for permanent brachytherapy. Both

Ra and

Rn have been replaced by newly developed isotopes for clinical brachytherapy, as discussed below.

3. Cesium-137:

Cs, a fi ssion by-product, is a po pular radium substitute because of its 30-year half-life. Its single J-ray (0.66 MeV) is less pen- etrating (HVL

=0.65 cm) than the J-rays from radium (HVL

=1.4 cm) or

Co (HVL

=1.1 cm).

Modern

Cs intracavitary tubes have been the

mainstay for intracavitary treatment of gyneco-

logical malignancies. The radioactive material

is distributed in insoluble glass microspheres,

which produce far less hazard from ruptured

sources than does the radon gas in a radium

tube. The active source material is then sealed in stainless steel encapsulation cylinders. Modern

137

Cs tubes usually have about a 2.65 mm exter- nal diameter with lengths of about 20 mm and active lengths between 14 and 20 mm, depend- ing on the vendor’s design. In addition, min- iaturized sources equivalent to 10 mg Ra with external diameters of about 1.25 mm attached to the end of long metal stems (Heyman-Simon sources) are used to treat endometrial cancer.

Cesium-137 needles were used as replacements for

Ra needles in interstitial implants; how- ever, their use has been diminishing in favor of more popular remote afterloading systems.

4. Cobalt-60:

Co: is produced from thermal neu- trons captured by

Co. The subsequent decay to

60

Ni releases two highly energetic J-rays (1.17 and 1.33 MeV), but

Co has a relatively short half-life (5.26 years). Cobalt-60 tubes and needles were used for brachytherapy during the 1960s and 1970s. Because of its high specifi c activity,

Co spherical pellets are used for HDR intracavitary therapy in some centers.

5. Iridium-192:

Ir, which has a 74-day half-life and lower-energy J-rays (average J-ray energy, 0.4 MeV), is the most widely used source for tem- porary interstitial implants. In Europe,

Ir is used in the form of a wire containing an iridium–

platinum radioactive core encased in a sheath of platinum. In the United States,

Ir is available as seeds (0.5-mm diameter by 3 mm long) with an active

Ir core cylinder contained in stainless steel or platinum encapsulation. The seeds are encapsulated in a 0.8-mm-diameter nylon ribbon and are usually spaced at 0.5 cm or 1 cm center- to-center intervals.

Ir ribbons and wires can be trimmed to the appropriate active length for each catheter. Finally, high-intensity

Ir sources are used in the latest-generation single stepping source HDR remote afterloading devices.

6. Gold-198: Insoluble

Au seeds, with a 2.7-day half-life and a 0.412 MeV J-ray, are available for use as a radon seed substitute to perform perma- nent implants. Gold-198 seeds are 2.5 mm long and 0.8 mm in outer diameter, and have 0.15- mm-thick platinum encapsulation.

12.4.2

Low-Energy Photon Emitters

The following information is given about low-energy photon emitters:

1. Iodine-125:

I seeds emit J-rays and X-rays with energies below 0.0355 MeV, has a half-life of 59.7 days, and are readily shielded by a few tenths of a millimeter of lead (HVL

=0.002 cm).

Many designs of

I seeds are currently avail- able, all having external dimensions similar to the Oncura model 6711 seed, with an outer cylin- drical encapsulation of titanium shell of 4.5 mm in length and 0.8 mm in diameter, as shown in Fig. 12.1. Iodine-125 seeds are used mostly for permanent implant treatments of cancers of the prostate, lung, sarcomas, as well as the tempo- rary implant treatment of ocular melanoma when affi xed in an eye plaque.

2. Palladium-103:

Pd, produced from thermal neutron capture in

Pd, is an alternative to

125

I for permanent implants.

Pd emits 20- to 23-keV characteristic X-rays and has a shorter half-life (16.9 vs 59.7 days). Because of the much higher dose rates at which

Pd doses are deliv- ered,

Pd is thought to have a greater biologi- cal effect than

I (Ling 1992; Ling et al. 1995;

Nath et al. 2005; Antipas et al. 2001; Wuu et al.

1996; Wuu and Zaider 1998). On the other hand, implants that use 103Pd may be more sensitive to errors in source positioning due to the reduced penetration of the low-energy x rays. These errors primarily affect the target tissue, and in cases such as prostate cancer implants where seed insertion errors are a few millimeters, the dose errors can be substantial (Dawson et al. 1994; Nath et al.

2000). Figure 12.2 shows a diagram of the Thera- genics model 200

Pd seed.

End weld Silver rod with ¹²⁵I

adsorbed to surface Titanium capsule 4.6 mm

0.8 mm

Fig. 12.1 Nycomed Amersham model 6711 ¹²⁵I seed. (From Rivard et al. 2004)

0.8 mm 4.5 mm

Titanium end cup

Titanium capsule

Lead marker Laser weld

both ends Graphite pellets

with ¹⁰³Pd coating

Fig 12.2 Theragenics model 200 ¹⁰³Pd seed. (From Rivard et al. 2004)

3. Cesium-131:

Cs has an average energy of 31 keV and a half-life of 9 days, and is thought to com- bine the advantages of higher energy of

I and the high dose rate of

Pd for permanent implants (Murphy et al. 2004; Yue et al. 2005). Cesium-131 seed sources, with external dimensions similar to

125

I and

Pd seed sources, have recently become available for permanent implant treatments.

12.4.3

Emerging Sources

Interest in use of ytterbium-169 as a brachytherapy sources dates back to the early 1990s (Mason et al.

1992; Perera et al.1994; Das et al. 1995). Ytterbium- 169 has an average energy of approximately 90 keV, with a half-life of 31 days. The average energy of

169

Yb falls within a region of Compton scattering interaction in tissue, where the ratio of scattered photon energy to primary photon energy is nearly at a maximum. The

Yb sources can therefore deliver a higher dose to points distant from the source in comparison with traditional brachytherapy sources, such as

Cs and

Ir, which is considered an advan- tage for gynecological cancer treatments. At the same time, radiation shielding for

Yb require- ments is much easier than for

Cs and

Ir, due to its lower energy and smaller half-value-layer value in lead (see Table 12.1).

12.5

Dose Calculations in Brachytherapy

12.5.1

The Superposition Principle

The clinical calculation of dose distribution from brachytherapy sources, as is currently practiced, is based on the superposition principle, i.e., the total dose distribution, at a given point of interest, from a group of brachytherapy sources is equal to the sum of the dose to that point by each of the brachytherapy sources in the group, or

where D

i

(x, y, z) is the dose contribution from the

source to point of interest (x,y,z).

The superposition principle assumes that the dose distribution to a point of interest is not

affected by the presence of other sources. In real- ity, this assumption is only an approximation.

The accuracy of this assumption, or the so-called interseed effect, depends on the average energy of the sources, as well as the distances of the points of interest to the sources. For low-energy sources, such as

I and

Pd seeds used in permanent pros- tate implants, this assumption has been shown to underestimate dose by several percent (Meigooni et al. 1992; DeMarco et al. 1995; Zhang et al. 2005;

Chibani et al. 2005). Similar effects have also been demonstrated for intravascular brachytherapy applications, where high-energy beta emitting sources are used to deliver a prescription dose to points located within 2 mm from the center of the sources, an extremely short distance in brachy- therapy applications (Patel et al. 2002). For high- energy photon emitting sources, such as

Cs and

192

Ir, the interseed effect is negligible.

Assuming that the superposition principle holds for a clinical application, the brachytherapy dose cal- culation problem reduces to the calculation of single sources. i.e., calculation of the radiation dose distri- bution around a single brachytherapy source. Once such dose distribution parameters are obtained, they can be tabulated for a manual calculation or for computerized isodose distribution calculation for an implant using a group of sources.

12.5.2

Source Strength Units

Brachytherapy source specifi cation protocols have evolved since its inception. The earliest unit for brachytherapy source strength was based on the

of Curie (Ci) for activity:

1 g radium = 1 Ci = 3.7u10

disintegrations/s.

While the unit of Ci, as defined in terms of ele-

mental disintegration rate, is a measurable physical

quantity, it cannot be easily applied to brachyther-

apy source strength specifications because the dose

distribution around an encapsulated brachytherapy

source depends on the attenuation and scattering of

the photons by the encapsulation material. Speci-

fication of source strength based on the elemental

disintegration rate is therefore usually referred to

as contained activity in brachytherapy literature,

and holds little interest to brachytherapy physicists

except in the case of regulatory compliances, where

the federal and state governments in the United States require accounting of radioactive material possession in terms of this quantity.

Brachytherapy source strength specifications, therefore, are usually based on what can be mea- sured outside of the encapsulated source. The fol- lowing units are often encountered in brachyther- apy literature:

1. Milligram-radium-equivalent (mgRaEq): High- energy brachytherapy sources with average energy higher than 300 keV have dose distribu- tion characteristics similar to that of radium.

They are usually referred to as radium substi- tute sources; 1 mgRaEq of the radium substitute source is defi ned to be the amount of the radium substitute source that gives the same output as a 1 mg radium source encapsulated in 0.5 mm platinum in the same output measurement geom- etry. The measurement geometry specifi cation includes a large distance between the source and the dosimeter, such that the radiation distribution is equivalent to a point source; that the dosim- eter should be placed on the transverse axis of the source; and air attenuation and scattering should be corrected. The output quantity of a brachytherapy source used in the determination of mgRaEq is exposure, with units in Roentgen (R). An amount of 1 mgRaEq of a radium substi- tute source therefore will have the same exposure as 1 mg of radium with 0.5-mm platinum encap- sulation, or 8.25 R cm

h.

The quantity mgRaEq has a long use history in clinical brachytherapy. The product mgRaEq and the implant time, mgRaEq h, has been used as a prescription quantity for many temporary implants, such as in tandem and ovoids implant for the treatment of cervical cancer.

2. Apparent activity (A): Apparent activity is defi ned similarly to mgRaEq, with the exception that the encapsulated radium source is replaced by an unshielded source of the specifi ed isotope and has the unit of Ci. A 1-Ci apparent activity of an encapsulated radioisotope source is defi ned to be the amount of encapsulated source that gives rise to the same output, or exposure in air, as an unencapsulated source of the same isotope of 1 Ci [contained] activity. Apparent activity, due to its not being based on radium sources, is applicable to non-radium-substitute sources such as

I and

103

Pd sources.

3. Air-kerma strength (S

): Both milligram-radium- equivalent and apparent activity in mCi have served the brachytherapy community for a long

time and hold historical signifi cance by their association with the clinical experiences accu- mulated over the years. They are both limited, however, in their applications, and are associated with historical variations in their conversion to exposure in air through the use of the exposure rate constant. In addition, for the calculation of dose in water, as is required for brachytherapy applications, an additional conversion factor between exposure in air and dose in water is required. The AAPM therefore recommended the use of air-kerma strength (S

), defi ned as the dose in free air along the transverse axis of an encap- sulated source, measured at a large distance from the source such that the source can be approxi- mated by a point source. Air-kerma strength has the unit of cGy cm

h

, and is represented by the symbol U.

12.5.3

Single Source Dosimetry

12.5.3.1

Point Source Dosimetry

The dose distribution surrounding a point brachy- therapy source will decrease with the square of the distance r from the source such that the dose rate

∝ 1/r

. Because the source strength specifications are defined for an output in air (exposure or air- kerma), a conversion factor from the quantity in air to dose in water is required. This is represented by the f

. The dose falloff is also affected by the attenuation and scattering of photons in media. When these factors are combined, the dose rate at a distance of r centimeters away from a point brachytherapy source is

.

This dose rate equation is appropriate for source

strengths specifi ed in apparent activity, where A is

the source activity, ( Γ

)

is the exposure rate con-

stant (converting the source strength to exposure

in air), with δ specifying the lower limit of photon

energy included in the determination of the expo-

sure rate constant (photons with energy lower than δ

are absorbed near source surface and do not contrib-

ute to doses at clinically significant target locations),

and x specifying the isotope. The factor f

has

units of cGy/R and is specific to photon energy. The

tissue attenuation and scatter factor, T(r), accounts for the attenuation and scattering of photons from the source as they traverse the medium.

12.5.3.2

Line Source Dosimetry

Clinical brachytherapy sources have finite physical dimensions, typically in the shape of a cylinder, and are encapsulated in a metal shell of stainless steel, platinum, or titanium. Dose calculations around such sources therefore must include considerations of the geometric distribution of the source within the encapsulated source, as well as the attenuation and scattering of the encapsulation materials.

Sievert Integral. In its simplest form, the Sievert

integral (Williamson et al. 1983; Williamson 1996; Karaiskos et al. 2000) only considers the effect of active source geometric distribution within the encapsulated source on the dose distribution around the source by integrating over the active source particles. For a source that can be approxi- mated by a line of length L and without encapsula- tion, the Sievert integral takes the form of

,

where r

is the vector between a segment dl on the line source and the point of interest P, as shown in Fig. 12.3.

When the attenuation and scattering of the pho- tons by the active source and the encapsulation materials are considered, as shown in Fig. 12.4, the dose at point P becomes

, where t1 and t2 are the thicknesses of active source and encapsulation materials along the vector r

to point P, and µ1 and µ2 represent the average linear attenuation coefficients for the average energy of the source photon spectrum in the active source and encapsulation materials, respectively.

The use of attenuation coefficients for the aver- age energy of the source photon spectrum in Sievert integral has been shown to be highly accurate with some high-energy sources such as Cs-137 tubes (Williamson 1996), allowing its implementation in commercial brachytherapy treatment-planning systems; however, for sources with complex photon energy spectra, such as Ir-192, and low-energy

sources, such as I-125 and Pd-103, Sievert integral results in significant dose calculation errors. Modi- fied forms of Sievert integral have been proposed to improve the accuracy of Sievert integral for Ir-192 sources (Williamson 1996; Karaiskos et al. 2000), although those are at the present time not available in commercial treatment-planning systems.

For high-energy sources with minimal active source and encapsulation thicknesses, the attenua- tion and scatter of photons in the source material and in the medium can be assumed to minimally affect the dose distribution. The terms of T( r

), e

, and

can then be removed from the integral. For

Fig. 12.3 Unfi ltered Sievert line source integral. The contribu- tions to dose at point P by each source segment dl are integrat- ed over the entire active source length L without considering the attenuation and scattering of active source materials and source encapsulation.

L

P

r dl

h = treatment depth

r

L dl

P

t1 t2

Source plane

Fig. 12.4 Integration over source segments dl through the en- tire active source length, taking into account of active source and source encapsulation attenuation by segments t1 and t2 in a Sievert Integral

a point of interest located on the transverse axis of the source, this yields the following unfiltered line source approximation of Sievert integral:

,

where x is the distance of the point of interest to the center of the source along its transverse axis. The unfiltered line source integral is often adequately accurate for certain clinical applications, such as in the manual calculation of dose at a point on the transverse axis of the source, for the purpose of double-checking a computer-generated dose distri- bution for a group of sources arranged in a line.

Away-Along Tables. While Sievert integral provides

adequate calculation accuracy for high energy sources, the integration cannot be quickly done manually. A look-up table summarizing the dose distribution surrounding a line source has been a powerful tool for quality assurance. A point in an away-along table is identifi ed by its distance away from the source projected to the transverse axis of the source, and its distance along the longitudinal axis of the source and projected to the source lon- gitudinal axis. Table 12.2 shows such a table for the 3M model 6500 Cs-137 source. Clinically, the away and along distances of a point of interest relative to

a source can be measured off a radiograph of the implant and the dose contribution of the source to this point looked up on the table. This process is then repeated for all sources in the implant to obtain the total dose to the point of interest.

TG43 Formalism. The line source dose calculation

formalisms discussed thus far, the Sievert integral and away-along tables, have served traditional brachytherapy dose calculation needs adequately for

226

Ra,

Cs, and

Au sources. Attempts have been made to apply these formalisms to newer brachy- therapy sources such as

Ir,

I, and

Pd seeds.

The use of Sievert integral for dose calculation of these sources has proven to be challenging, due to this source’s complex photon emission spectrum (Williamson 1996; Karaiskos et al. 2000), while computational errors increase rapidly due to inter- polation for points near the sources and between the away-along table entries. In 1995, the American Association of Physicists in Medicine (AAPM) Radi- ation Therapy Committee Task Group 43 published its report (Nath et al. 1995) entitled “Dosimetry of Interstitial Brachytherapy Sources,” subsequently revised in 2004 in the updated TG43 report (Rivard et al. 2004). This report introduced a dose calcu- lation formalism, commonly referred as the TG43 formalism. The TG43 formalism is based on using

Table 12.2 Away-along dose distribution table for 3M model 6500 Cs-137 source. Unit of the entries is cGy × h^–1/(µGy × m²× h^–1) source strength. (From Williamson 1998a)

Distance along (cm)

Distance away (cm)

0.00 0.25 0.50 0.75 1.00 1.50 2.00 2.50 3.00 3.50 4.00 5.00 6.00 7.00 7.00 0.0193 0.0189 0.0184 0.0180 0.0179 0.0178 0.0176 0.0170 0.0164 0.0156 0.0147 0.0129 0.0111 0.0094 6.00 0.0269 0.0263 0.0254 0.0249 0.0248 0.0247 0.0241 0.0231 0.0218 0.0204 0.0189 0.0160 0.0134 0.0112 5.00 0.0397 0.0386 0.0370 0.0365 0.0365 0.0359 0.0344 0.0322 0.0297 0.0272 0.0247 0.0201 0.0162 0.0131 4.00 0.0638 0.0614 0.0586 0.0584 0.0582 0.0559 0.0517 0.0468 0.0416 0.0367 0.0323 0.0249 0.0193 0.0151 3.50 0.0848 0.0811 0.0774 0.0773 0.0766 0.0719 0.0648 0.0570 0.0495 0.0428 0.0369 0.0276 0.0209 0.0162 3.00 0.118 0.112 0.107 0.107 0.105 0.0949 0.0824 0.0700 0.0591 0.0497 0.0420 0.0304 0.0226 0.0172 2.50 0.176 0.164 0.159 0.156 0.149 0.128 0.106 0.0863 0.0702 0.0575 0.0475 0.0332 0.0241 0.0181 2.00 0.290 0.265 0.257 0.246 0.225 0.178 0.137 0.106 0.0827 0.0657 0.0530 0.0359 0.0257 0.0189 1.50 0.580 0.516 0.489 0.427 0.360 0.249 0.176 0.128 0.0957 0.0737 0.0581 0.0383 0.0268 0.0196 1.00 – 1.580 1.135 0.799 0.582 0.34 0.217 0.149 0.107 0.0807 0.0625 0.0403 0.0278 0.0202 0.50 – 6.569 2.468 1.345 0.852 0.426 0.252 0.165 0.116 0.0855 0.0654 0.0415 0.0284 0.0205 0.00 – 7.806 3.039 1.594 0.973 0.462 0.266 0.171 0.119 0.0872 0.0664 0.042 0.0286 0.0206 –0.50 – 6.566 2.466 1.343 0.851 0.425 0.252 0.165 0.116 0.0855 0.0654 0.0416 0.0285 0.0205 –1.00 – 1.590 1.136 0.803 0.584 0.340 0.217 0.149 0.108 0.0807 0.0625 0.0403 0.0278 0.0202 –1.50 0.547 0.498 0.489 0.428 0.360 0.249 0.176 0.128 0.0958 0.0738 0.0582 0.0384 0.0269 0.0196 –2.00 0.273 0.251 0.256 0.247 0.226 0.178 0.137 0.106 0.0828 0.0657 0.0530 0.0360 0.0256 0.0189 –2.50 0.166 0.154 0.155 0.156 0.149 0.129 0.106 0.0863 0.0702 0.0575 0.0475 0.0333 0.0241 0.0181 –3.00 0.112 0.106 0.104 0.106 0.104 0.0949 0.0824 0.0701 0.0591 0.0497 0.0420 0.030393 0.0226 0.0172 –3.50 0.0802 0.0767 0.0745 0.0759 0.0760 0.0719 0.0648 0.0571 0.0495 0.0428 0.0369 0.027559 0.0209 0.0162 –4.00 0.0604 0.0582 0.0561 0.0570 0.0575 0.0557 0.0517 0.0468 0.0416 0.0367 0.0322 0.024838 0.0193 0.0151 –5.00 0.0376 0.0366 0.0352 0.0353 0.0358 0.0357 0.0344 0.0323 0.0298 0.0271 0.0246 0.019982 0.0161 0.0131 –6.00 0.0255 0.0250 0.0242 0.0239 0.0242 0.0244 0.0240 0.0231 0.0218 0.0204 0.0189 0.016003 0.0134 0.0111 –7.00 0.0183 0.018 0.0175 0.0172 0.0173 0.0175 0.0174 0.0170 0.0164 0.0156 0.0147 0.012847 0.0111 0.0094

air-kerma strength for source strength specifi cation and is described by the following equation:

,

using the coordinate system shown in Fig. 12.5, where S

is the source strength specified in air-kerma strength, in units of U=cGy cm

h

. Clinically used brachytherapy sources should have their strength directly or secondarily traceable to a calibration standard established by the National Institute of Standards and Technology (NIST).

Λ is the dose rate constant of the source in water, in unit of cGy h

U

. It is defined to the dose rate at 1 cm from a source of 1 U strength, along the source transverse axis, give by Λ = D(r

, θ

)/S

. The dose rate constant for a given source must be evaluated carefully, using either well-validated calculation methods, such as Monte Carlo calculations, or mea- sured using appropriate dosimeters, such as ther- moluminescent dosimeters (TLD).

G_L(r,

θ) is the geometry function that describes the effect of the active source material distribution within the source on the dose distribution out- side the source, and is by the inverse square law

culated by integrating the inverse-square law over all active source particles within an encapsulated brachytherapy source. In practice, it is common to approximate the active source material distribu- tion within a brachytherapy source by an idealized geometry such as a line. The values of the geom- etry function can then be analytically calculated, thereby avoiding interpolation errors at short dis- tances to the source as may occur with the use of away-along tables. It is, however, crucial that the assumptions made in calculating the geometry function, such as the length of the idealized source distribution be consistent between the source dosimetry parameter derivation and the clinical applications of these parameters. Disagreement in the values of these assumptions may lead to sig- nifi cant dose calculation errors at points close to the source. The updated TG43 report emphasizes this point by using a subscript L in the symbol for the geometry function, indicating the use of a line source assumption for the calculation of geometry function values.

g(r), the radial dose function, accounts for the

effect of photon absorption and scattering on the dose distribution along the source transverse axis.

F(r,

θ) is the anisotropy function, which describes the effect of anisotropic photon attenuation, either

by the source materials (active source core and source encapsulation) or at locations away from the source transverse axis.

12.5.3.3

Total Delivered Dose Calculations

Given the half-life value of a radioactive isotope and the initial dose rate D

0

(r) at point r from the source, the total dose delivered to point r in the time interval [0, t

] can be calculated to be

, using the relation given in Eq. (2a).

For short treatments using sources with large half-lives, the term

can be adequately approximated by (1 – ln(2) × t

/ T

), resulting in D(r) = D

0

(r) × t

. This assumption, however, does not apply to some brachytherapy treatments using short half-lived isotopes, such as

I or

Pd seeds. Of particular interest is the use of these sources for permanent implants, where the total doses delivered to a point of interest is given by

D(r) = D.

0

(r) × 1.433 × T

,

Because of the special importance of the term 1.433 × T

, it is defined to be the average life of an isotope, i.e., T

= 1.433 × T

.

t

P(r_,θ0)

r_ = 1cm θ

P(r,θ)

z L

r x

θ1 θ2

β

Fig. 12.5 Coordinate system of TG43 dose calculation formal- ism. (From Rivard et al. 2004)

12.6

Gynecological Intracavitary Implant

12.6.1 Applicators

Brachytherapy implants are an integral part of the treatment of many gynecological cancers, includ- ing the cervix, uterine body, and vagina. Applicators are used to hold the brachytherapy sources in clini- cally defined configurations, or loading patterns.

The applicators used for cervical and uterine cancer treatments typically include a tandem, to be inserted into the uterus, and two ovoids, to be positioned in the vaginal vault abutting the cervix. The Fletcher- Suit-Delclos applicator is one of such applicator sets commonly in clinical use in the United States (see Fig. 12.6). This applicator set has tandems of sev- eral curvatures to conform to the patient’s anatomy, as well as ovoids of diameters of 2, 2.5, and 3 cm, with the larger diameters achieved by fitting plastic caps outside the 2-cm-diameter ovoids. Internally, the ovoids have tungsten shields in the anterior and posterior aspects of the ovoids, as shown in Fig. 12.7, to provide dose attenuation and reduced doses to the bladder and rectum.

12.6.2

Dose Specification for Cervical Cancer Treatments

In the United States, cervix cancer treatments using brachytherapy implants are typically prescribed by one of two methods: the total exposure method, as represented by the product of the total strength of sources implanted and the total source dwell time; and the point-A prescription method. The total exposure method is used at Washington University at St. Louis, where a typical prescription for a course of low-dose rate cervical cancer treatment includes two insertions to deliver nominally 8000 mgRaEq h exposure, with the actual delivered exposure modified based on the length of the tandem and the diameters of the ovoids (Williamson 1998b). It is noteworthy that, when pre- scription by total exposure is chosen for a treatment, the pattern of source loading, or the distribution of source strengths in the tandem and ovoids applica- tors, should adhere to institutional rules. Williamson (1998b) explained these rules in detail.

The traditional Manchester system for cervical cancer brachytherapy specifies the prescription dose at point A, defined to be the paracervical points at

2 cm superior to the vaginal fornix, and 2 cm lateral from the uterine canal. The system also specifies point B, located at the same superior–inferior level as point A, but at 5 cm lateral from the patient’s mid- line, intended to represent dose to the parametria. It is noted that point A is related to the orientation of the cervical canal, as localized by the tandem in a radiograph. Lateral distension of the cervical canal results in the corresponding shift of point A.

12.6.3

ICRU Report 38 Recommendations

The International Commission on Radiation Units and Measurements (ICRU) made several recommen- dations (ICRU 1985) on the reporting of cervix cancer brachytherapy treatment dosimetry, including the volume included by the 60 Gy isodose line, estimated by the product of the length, width, and height of this isodose line. In addition, ICRU report 38 clarified the reporting of bladder and rectum doses. The bladder dose is measured at the posterior-most aspect of a 7-cc Foley balloon in the bladder, pulled back such that the balloon is located on the bladder trigone.

The rectum point is defined by the point bisecting the ovoid sources supero-inferiorly, and at 5 mm poste- rior to the posterior vaginal wall. Figure 12.8 shows the ICRU report 38 definitions of the bladder and rectal points, together with the Manchester system point-A and point-B definitions.

12.6.4

Volumetric Image-Based GYN Brachytherapy The treatment planning of GYN tandem and ovoids, to date, are still often based on planar images,

Fig. 12.6 Fletcher-Suit low-dose-rate cervix applicator set (Best Medical, Springfi eld, Virginia)

unable to reap the benefits of volumetric images in target and critical organ definitions. This lack of progress has largely been due to the lack of shielded ovoid applicators that are free of CT and MR imag- ing artifacts. Conventional ovoid applicators, with their tungsten rectal and bladder shields, create so many imaging artifacts that accurate critical organ and target segmentation are impossible. In addi- tion, the progress of volumetric image-based GYN brachytherapy has been limited by CT’s low specific- ity for delineating abdominal tumors.

Weeks and Montana (1997) developed the first CT-compatible ovoid applicators with afterloadable shields. The rest of the applicators are made of alu- minum, producing little CT imaging artifacts. The CT images of the implants using these applicators may be acquired first, before the shields and the sources are inserted. The applicators have large- diameter handles for the shields to pass through.

Martel and Narayana (1998) used these appli- cators to perform 3D treatment planning of GYN tandem and ovoids implants, and obtained the first set of 3D dose distribution data, especially for the rectum and bladder, to show that the maxi-

mum doses that these critical organs receive in a tandem and ovoid implant are significantly higher than estimated by the ICRU report 38 rectal and bladder points. Commercial CT and MR compat- ible HDR tandem and ovoids, made of carbon fiber or titanium, are now available for use in volumet- ric image-based GYB tandem and ovoid implants, although those still suffer from the absence of high-density rectal and bladder shields, making it difficult to translate the large amount of clini- cal experience in GYN tandem and ovoid implants, established using applicators with rectal and blad- der shields, into the implementation of this new technology.

Magnetic resonance imaging has been used recently for the target and critical organ delineation in tandem and ovoid implants. Compared with CT, MR imaging provides significantly higher specificity for tumor delineation, while preserving the ability to allow critical organ segmentation. For these reasons, the ABS has recommended its use for image-based GYN brachytherapy (Nag et al. 2004). Modifying the International Commission on Radiation Units and Measurements (ICRU) report 50 nomenclature

S

Sec D - D

Shielding Details

A

.SS

20 D|A Sec B - B

A

1.3 SS C

D D

.7 SS

C

B B

3C

Sec A - A Sec C - C

5 Tungsten (92%)

1. SS

Fig. 12.7 Internal structures of Fletch- er-Suit-Delclos ovoid applicator (Best Medical, Springfi eld, Virginia)

(ICRU 1993), the ABS defines the following target volumes for image-based GYN brachytherapy:

1. GTV

: tumor visible in imaging

2. GTV: GTV

+ clinically visible or palpable tumor (including parametria)

3. GTV+cx: GTV plus the entire cervix 4. pCTV (primary CTV)

a. For external beam, GTV+cx; entire uterus;

parametria; and upper 2 cm of vagina b. For brachytherapy, GTV+cx plus 1-cm margin 5. rCTV (regional CTV) for external beam: pCTV

plus regional lymph nodes

6. CTV for external beam: combination of pCTV for external beam and rCTV

Note that, in the above, separate target volumes have been defined for treatment planning using external beam techniques and brachytherapy treat- ments, respectively. The ABS further recommends that individual MR imaging sessions be performed prior to the initiation of external therapy treat- ments, and before each fraction of brachytherapy treatments following the insertion of brachytherapy applicators, as the tumor regresses through therapy, and as its shape and location are affected by the insertion.

Positron emission tomography (PET) has recently received increased attention for use in GYN brachy- therapy (Mutic et al. 2002; Malyapa et al. 2002;

Wahab et al. 2004; Lin et al. 2005). The PET imag- ing has the potential to identify biologically active tumor regions, and, when co-registered with CT images, allows for biologically determined tumor delineation and anatomically segmented critical organs. The PET images, acquired prior to each brachytherapy fraction with the applicators inserted, demonstrate clearly tumor regression through the course of radiotherapy treatment, and allow intri- cate sculpting of the tumor, as represented by high- uptake regions on the PET images for planning of optimized HDR brachytherapy treatments.

12.7

Interstitial Implant Dosimetry Systems Brachytherapy treatments using interstitial tech- nique insert brachytherapy sources within the target volume, in order to deliver a prescribed target dose with acceptable dose distribution homogeneity.

Prior to the development of computerized treatment- planning techniques, several classical implant sys- tems were developed to calculate, for a given target volume, the total activity of the sources, number of sources, and the source distribution within the target volume, for a given prescription dose. The relation between the target dimensions and the total activity were given in tabular form for a nominal prescription dose, whereas the rules of source distri- bution were specified separately. While the impor- tance of these classical systems has been reduced with the use of computerized treatment planning, they remain fundamental in the planning of inter- stitial brachytherapy treatments, both to help guide the pattern of source distribution within the target volume for improved dose distribution homoge- neity, and to ensure the technical consistency of

2 cm

6 cm

5 cm

2 cm

A A B P

Uterine sources

Ovoid source

Ovoid source Spacer

•

6 cm 5 cm

2 cm 2 cm

Patient midline

Patient transverse plane A

A

B P

B

Fig. 12.8 ICRU report 38 defi nition of bladder and rectal points and Manchester system point A and point B (ICRU 1985)

treatment delivery for all patients. In addition, the classical implant systems are often used as tools of independent quality assurance checks of the com- puter treatment plans.

12.7.1

Paterson–Parker (Manchester) System

The Paterson–Parker system, developed by Paterson and Parker in 1934, aims to deliver a uniform dose (±10% from the prescribed or stated dose) on the plane or surface of treated volume. The sources are distrib- uted non-uniformly following specific rules, based on the size of the target volume, with more source strength concentrated in the periphery of the target volume. Such non-uniform distribution of source activities may be achieved either by use of sources of non-uniform strengths, or by varying the spacing of sources of uniform strengths. The Patterson–Parker dose tables give the cumulative source strength required to deliver 860 cGy, using current factors and dose units, as a function of the treated area (planar implants) or volume (Williamson 1998b).

Regarding single-plane implants, source cathe- ters, arranged in a single plane at 1-cm spacing, can be used to deliver a prescribed dose to a target plane at 0.5 cm away and parallel to the source plane, as shown in Figure 12.9. Cross-end needles may be used, in which case the lengths of the needles in the plane may be reduced by 10% for each cross-end needle.

The fractions of source strengths in the periphery of the implant depends on the total treated area: for areas less than 25 cm

, two-thirds of the total activi- ties are implanted in the periphery; changing to one-half for total areas between 25 and 100 cm

and one-third for total areas larger than 100 cm

, respec- tively. The total activity is further increased for non- rectangular target areas. For thicker slabs of target, up to 2.5 cm thick, the needles may be arranged in two parallel planes. The total activity needed for double-plane implants are looked up using a single plane implant table, followed by application of a correction factor depending on the thickness of the target tissue. The total activities are then distributed between the two source planes in proportion to their relative areas.

Regarding volume implants, the Patterson–

Parker system for volume implant is similar to the planar implant, in that the total activity is non-uni- formly distributed between the periphery and the core (center) of the target volume. Typically, with all sides of a target volume implanted, the total activity

is divided into eight parts, with only two parts of the total eight parts in the core volume of the target.

The lengths of the implant needles may be reduced if cross-end needles are used, by 7.5% for each end of the volume implanted with cross-end needles.

12.7.2

Quimby System

This system, developed by Quimby in 1932, is based on a uniform distribution of source strength, allowing a higher dose in the center of the treat- ment volume than near the periphery. Typically, for equal dose delivery to similar size planar or volume implants, the total source strength required when using the Quimby system will be much greater than what is required by the Patterson–Parker system.

12.7.3 Paris System

The Paris system is used primarily for single- and double-plane implants using parallel, equidistant needles arranged in triangular or rectangular shapes when viewed on the needles’ ends. All sources used in a Paris system-based implant are to have the same linear strength, although it is possible that sources of different lengths may be used. A central implant plane is defined that approximately bisects all implanted parallel needles. The prescribed dose is a percentage, typically 85%, of the average doses at minimum dose points within the central plane, or the so-called basal dose points. Geometrically, these minimum dose points are approximated by points equidistant to neighboring needles. The lengths and spacing of the needles are therefore depen- dent on the target thickness and width. Due to the

h = Treatment depth

Source plane Treatment plane

Fig. 12.9 Planar implant of Paterson-Parker system