13 Modern Principles of Brachytherapy Physics

From 2-D to 3-D to Dynamic Planning and Delivery

Brachytherapy is now a century old. It has grown through many generations and evolved into a very efficient and uniform way to deliver radiation to tumors. A complete week, with approximately 30 hours of lectures, was devoted to cover physics and the technical aspects of brachytherapy at the American Association of Physicists in Medicine (AAPM) Summer School in 2005. The book ¹ about this event certainly represents one of the most comprehensive reviews of the field, and many textbooks and thousands of scientific papers have also been devoted to this topic. Nevertheless, the majority of the concepts reported in the literature were developed in a different era, when three-dimensional (3-D) imaging was limited and dose planning was based on “systems” that tried to prevent hot spots and suboptimal dose distributions in the absence of other safeguards. Clearly, many of the paradigms of brachytherapy deserve to be revisited.

Programmable remote afterloaders, new radiation and electronic sources, more accurate dose calculation algorithms, 3-D imaging modalities (computed tomography [CT], magnetic resonance imaging [MRI], positron emission tomography [PET]-CT, ultrasound [US], etc.), explicit anatomy-based dose optimization planning tools (inverse planning), as well as image guidance and robotic dose delivery lead the way to modern brachytherapy. These changes are all happening at once. This chapter cannot pretend to present a complete review of the new trends. They are growing in many directions and being refined through scientific experiments and clinical protocols. This paper can merely present a subjective snapshot of where the field is and where it is going. If this chapter can raise the confidence level and contribute to accelerating the transition to modern brachytherapy, then our objective will have been fulfilled.

Why 3-D in Brachytherapy?

There are several reasons that support the transition from planning being performed on x-ray films to planning using a 3-D image data set. Some of these reasons are intuitive and practical. Others are based on scientific experiments or clinical experience. One could certainly learn from history. External-beam radiation therapy, although far more recent than brachytherapy, has now completely transitioned from two-dimensional (2-D) to 3-D–based planning.

Whether it is a permanent implant, pulse-dose-rate (PDR) or high-dose-rate (HDR), brachytherapy is a high-dose procedure. In the case of permanent implants, the dose rate is low and the dose is delivered over several months. However, the position of the seeds is determined during one insertion and the entire dose is delivered in this configuration. From this point of view, even permanent implants are high-dose procedures. All other high-dose procedures (Cyberknife, gamma knife, stereotatic radiosurgery, etc.) include complete and explicit 3-D imaging for planning (either MRI or CT, or both) and a rigid reference frame.

Because brachytherapy has the potential to deliver highly conformal dose distribution by fully exploiting the inverse square law and the source distribution, it can best benefit from the 3-D information. In fact, although computer-controlled afterloaders have been available for more than 15 years to deliver complex dose distributions, only 3-D dose planning can fully exploit their capabilities.

Inherently, 3-D images provide more anatomic information. Most soft-tissue is not visible on x-ray films. The tumor definition, its relation to the surrounding anatomy and the organs at risk (OARs), and the position of the brachytherapy applicator or catheters relative to the anatomy are better assessed with 3-D imaging. When sources are placed, or dwell times adjusted, the dose distribution can be calculated and the user can assess the amount of dose received by the organs, also opening the possibilities to compute organ-specific dose-volume histograms (DVH). Effectively, dose distributions are more accurately evaluated using DVH and derived dosimetric indices available only on 3-D planning systems in which target and organs can be delineated and dose distributions calculated in three-dimensionally. Furthermore, when multi-imaging modalities are used (CT, MRI, PET-CT, etc.), it is only through image registration and fusion that one can achieve an accurate transfer of information. A prerequisite to using the information is the availability of 3-D–based planning.

Many treatment centers have reported the practical and clinical benefits of 3-D planning after transitioning from 2-D to 3-D. As demonstrated by several studies, in the case of gynecologic treatments, much higher bladder and rectal doses are found when using the CT information. The minimum dose (D^min) to the cervix target volume is lower than the dose to point A in each case. Similarly, dose to normal tissue is generally reduced and smaller volume is irradiated when 3-D planning is introduced. Image fusion of CT and MRI data enabled improved target volume definition in 3-D–based planning.²

CT-based planning allows for some automation in the anatomy delineation and catheter definition. For instance, automatic catheter reconstruction in prostate HDR treatment can easily be achieved.³ The time required for the catheter reconstruction process using an autoreconstruction method is significantly reduced compared with the manual approach on CT, and is reduced even more when compared with film-based reconstruction.

Furthermore, dose escalation could hardly be considered if 3-D planning is not the standard. Several studies have shown that the dose delivered to single points do not correlate with side effects. For quality assurance, it is safer to review the entire 3-D volume. Each patient is unique. Each site and organ has its own dose “constraint.” A truly personalized treatment clearly requires an anatomy-based, dose-optimization planning approach. For dosimetry, the new era is inverse planning. For a long time, we have convinced ourselves that it is sufficient to look at dwell times or how the seeds are distributed to understand dose distribution. But the objective of radiation therapy is to deliver a specific dose distribution, not to achieve a specific seed configuration. The dose delivered is the central point, not the dwell-time or seed distribution per se.

We have all the motivations and evidences to be convinced that brachytherapy must be planned and delivered in 3-D. This is the philosophy followed everywhere in this chapter. This implies that every single step is performed using the complete 3-D data of patient anatomy, applicators, and catheters, as well as dose optimization and dose distribution evaluation.

Radiation Sources and Dose Calculation

This section is not intended to replace a radiation physics course. For that purpose, we refer the reader to an excellent book by Baltas et al.,⁴ which goes into much more detail or the handbook published by Mayles et al.⁵ We will, however, recall the basic concepts necessary to introduce modern brachytherapy radiation sources and dosimetric concepts. In particular, this should help the reader to relate known quantities such as kerma and dose to the AAPM Task Group 43 (TG43) formalism.⁶ Note that in this chapter, our interest is with photon-emitting sources.

Radioactivity and Interactions

Radioactivity defines the process by which unstable nuclei or isotopes undergo decay (with one or more transitions) that lead to stable nuclei species. The number of decays of a radioactive isotope within a finite period of time is given by:

Eq 1

Thus, after a period of time t, the remaining number of atoms, N_t, is given by:

Eq 2

where λ is the radioactive decay constant (probability to undergo a nuclear transformation per unit of time) and N₀ the initial number of atoms. The half-life, corresponds to the time necessary for one half of the initial number of atoms to decay or N = N₀/2. It also corresponds to half of the initial activity. Solving equation 1, one finds:

Eq 3

As for the decay constant, the half-life is specific to a given nucleus. The activity, A, is the number of disintegration per second:

Eq 4

Defining A₀ as λ N₀, one can also predict the activity at time t:

Eq 5

The international system (SI) unit for activity is the Becquerel (Bq) (one disintegration per second). The older unit, the Curie (Ci), corresponds to 3.7 × 10¹⁰ Bq. Note that historically, 1 Ci was defined as the activity of 1 g of radium. However, Marie Curie’s gram of radium was not exactly 1 g; thus the accepted value of 1 g of radium is 0.95 Ci. Finally, a related quantity of interest is the amount of activity per unit of mass (Bq/kg) or the specific activity:

Eq 6

where N_A is the Avogadro constant and M is molar mass. The right term of that equation applies to a pure isotope.

Let us now consider the fate of the products of the decay of a radioactive source. As the emitted photons travel in a given medium, interactions can occur. Among the possible results of an interaction are the scattering of the incident photons or partial loss of energy of the incident photons and complete absorption in the medium. The probability that an incident photon within a unit area will interact with an interaction site of the medium is given by the cross section σ (cm² or barn with 1b = 10⁻²⁴ cm²). The larger the cross section or the number of interaction sites, the more likely the interaction is to occur. Consider a source providing a fluence Φ (number of particle per unit area); the expected number of interaction per interaction site is given by:

Eq 7

Thus the lost of fluence dΦ or the variation dN in the number of incident particle N over a distance dx can be written as:

Eq 8

This is a well-known differential equation, the solution of which describes the exponential attenuation of the primary fluence; µ is the linear attenuation coefficient or the probability of interaction per unit of path length. For a given medium, the number of interaction sites per unit volume depends on the mass density ρ. To remove this dependency, the mass attenuation coefficient µ/ρ is used

Eq 9

in which A is the atomic mass number and N_A is the Avogadro constant. For photon sources, the relevant interactions are Compton, photoelectric, Rayleigh, and pair production, although for most modern brachytherapy sources, this last process is negligible. Each one has its own cross-section, which are combined to give the linear or mass attenuation coefficients (total cross section).

Dosimetry

For dosimetric purposes, we are of course interested by the energy deposited per unit of mass or dose (J/Kg or Gy). However, contrary to charged particles, photons must interact to impart energy to a medium. It is thus the secondary charged particles, electrons in the case of interest, that contribute to the dose. Furthermore, when a photon interacts, a (variable) fraction of its energy can be transferred to a secondary particle and spent in the medium and the remaining fraction is carried away from the interaction site by the scattered photon. The kinetic energy per unit of mass transferred to charged particles by uncharged particles is called kerma (K). The kerma can be related to the basic source characteristics such as photon fluence and energy. For a monoenergetic photon source, the kerma can be expressed as:

Eq 10

in which dE_tr is the average transferred energy, µ_tr/ρ is the mass energy transferred coefficient and ψ is the energy fluence (related to the photon fluence as previously defined). Note that the mass energy transferred coefficient is related to the mass attenuation coefficient through the fraction of the total energy (E) that is transferred per unit of path length.

Eq 11

Of course, the values of both coefficients depend on the energy of the interacting photon and the composition of the interacting medium. We can be even more specific and ask that the energy not only be transferred but also spent in collisions (as opposed to radiative loss by bremsstrahlung). In this case, the quantity of interest is the collision kerma (K^coll):

Eq 12

in which µ_en/ρ (or µ_ab/ρ) is the mass energy absorption coefficient and g is the mean fraction of energy not spent in collision with orbital electrons. In the context of charge particle equilibrium, the dose (D) is equal to the K^coll and thus can be related to the photon fluence and energy fluence.

Eq 13

Note that for materials naturally found in the human body, µ_tr/ρ is equal to µ_en/ρ to within 1% for photon energy up to 1 MeV. Thus, for most modern brachytherapy sources, radiative loss is negligible, K ≈ K^coll, and the last two expressions are basically equivalent. It is important to remember that the dose in a given medium can be related to the dose in a second medium, assuming no change in fluence and charged particle electronic equilibrium, by the following equation:

Eq 14

in which f_as is a factor correcting for attenuation and scattering when the air around the equilibrium mass is replaced by water (adopting the notion of Baltas et al.⁴). In a more compact form, this equation can be written as:

Eq 15

Average mass energy absorption coefficient can also be used based on the effective source energy. Note that the ratio of the mass energy absorption coefficient is sometime referred to as the factor (f_med).⁷

Source Specification and Air Kerma Strength

Source specification or strength has been expressed by various quantities: equivalent mass of radium, activity (in Ci), apparent activity for sealed source (i.e., activity equivalent to an unsealed source in terms of exposure rate), exposure rate constant, and so on. In modern dosimetry protocol, the AAPM recommends the use of air-kerma strength, S_k, to characterize the source strength. As its name indicates, it uses a measure of the collision kerma in air. Although it would be practical to express a source strength in terms of dose rate in water, this is very difficult for a low-energy brachytherapy source because of large dose gradients and extremely small temperature variation resulting from dose deposition. However, for these energies, exposure rate measurements in air are well understood. The exposure (or exposure rate) is related to the collision kerma as follows:

Eq 16

In which W_air/e is 33.97 J/C. The SI unit for exposure (exposure rate) is C Kg⁻¹ (C Kg⁻¹ s⁻¹). Here we have replaced the collision air kerma by the air kerma by neglecting the (1 − g) factor, as discussed before. Furthermore, exposure rate and air-kerma rate are more practical to describe the energy imparted by a radioactive source used because of the activity decay with time. Note that the air-kerma rate will vary as a function of distance to the source because of the inverse square law and attenuation in air. The latter is, however, negligible, leaving the geometric factor the dominant one. Therefore, the product of the air-kerma rate with the square of the distance should be roughly constant over a range of distances that are useful in clinical practice. The air-kerma strength, S_k, has been adopted first as a reference quantity for low-energy sources used in permanent seed implants (PSIs). It is defined as the product air-kerma rate in free space at the distance of measurement and the square of that distance along the perpendicular bisector

Eq 17

The distance r must be large enough compared with the extent of the source so that it can be treated as a point source. The indices δ in the air-kerma rate was included in the revised TG43 publication ⁸ to remove very low-energy photons produced close to the surface of the capsule or source cladding. These do not contribute to the dose beyond the first millimeter. The use of δ is similar to that of the air-kerma rate constant, γ_δ, as defined by the International Commission on Radiation Units and Measurements (ICRU). Values of γ_δ for the most commonly used isotopes can be found here.⁹ The SI unit for the air-kerma strength is cGycm²h⁻¹ (µGym²h⁻¹). The symbol U is also commonly used in practice and publications, but it is not an SI unit. At a distance (r) of 1 cm, the air-kerma strength of a perfect point source will give a dose rate in cGy/h in air. The air-kerma strength can be related to the apparent activity of a source by the air-kerma rate constant (we are not advocating the use of apparent activity):

Eq 18

IN which γ_δ has a value of 1.27 cGy cm² mCi⁻¹ h⁻¹ for ¹²⁵I, 1.30 cGy cm² mCi⁻¹ h⁻¹ for ¹⁰³Pd, and 4.205 cGy cm² mCi−¹ h⁻¹ for ¹⁹²Ir.

From the derivations we have conducted previously, we know how to convert a dose in air to a dose in water, so combining all the related equations, the dose in water for an ideal point source is given by

Eq 19

Let us now consider the dose rate at a reference point. In TG43, this point is defined at r₀ = 1 cm and θ = 90°:

Eq 20

Λ is referred to as the dose rate constant. It is defined for a given medium, water in this case, at the reference point. Its units are cGy h⁻¹ U⁻¹. It can be obtained by measuring directly the dose at the reference point in water and either divided by the air-kerma strength or calculated by the Monte Carlo method. The dose rate constant depends on the isotope and the source design, including the encapsulation.

Finally, let us consider the case in which the effective treatment time is long compared with the source half-life (such as in PSIs). The source strength will decay with time according to:

Eq 21

For a PSI, the integration from t = 0 to infinity of equation 21 is a constant related to the source half-life, and the total deposited dose is given by:

Eq 22

Revised TG43 Formulation

Up to now, we have not addressed the particularity of the various seed designs and how it affects the dose at other points than the reference point for a “prefect” point source. The TG43 formalism is a modular analytical equation in which the effect source geometry, attenuation, scattering, and anisotropy in the medium of interest are represented by different terms and accepted values (measured or calculated) are tabulated. Because these parameters are obtained for each source individually, the formalism can be used for any encapsulated sources, within the limit of the approximations stated previously. The following equation gives the dose rate at any point (r, θ), in polar coordinates:

Eq 23

In which X is replaced by “L” for a line source and “P” for the point source approximation. Let us define the various parts of this equation. We refer the reader the AAPM task group report or to more physics-oriented textbooks for a complete derivation of these terms. Note that the product of g_X(r) · F(r, θ) plays the same role as f_as in our previous derivation. S_k and Λ were defined previously.

G(r, θ) represent the geometry function (i.e., the variation of the radiation field intensity with distance compared with the reference point [r₀, θ₀] already defined). For a perfect point source, there will be complete symmetry in θ and thus G(r) = 1/r², such that

Eq 24

Obviously, at the reference point, this ratio is 1 by definition for both one-dimensional (1-D) and 2-D formalisms. However, for a cylindrical source of finite active length, assuming an ideal line for which the activity is uniformly distributed along the length of the source length, one needs to integrate contribution of each infinitesimal element. The geometry factor for a line (2-D) source is expressed as:

Eq 25

L_s is the active source length and β corresponds to the angle defined by the point of interest (r, θ) and the two ends of the active portion of the source. L_s is specific to the source design and is generally smaller than the physical length of the source. The reader should be reminded that the term at θ = 0° is new to the revised TG43 formalism. It handles the obvious problem of a division by 0 in the second term. However, at this time, most commercially available software only use the second term.

The radial dose function or g_X(r) describes the dependence of the attenuation and scatter of the radiation in the water medium as a function of the distance r on the transverse axis (i.e. θ = 90°):

Eq 26

By definition, equation 26 results in g_X(r = r₀) = 1. Thus, the variation in the radial dose function is given relative to that reference point. The removal from this equation of the geometric dependence and any variation in θ also has to be taken into account, as we describe later in this chapter. The radial dose function g_X(r) is obtained as function of r from measurements or more commonly by the Monte Carlo calculation. The 1-D or 2-D values are extracted using the corresponding point source or line source geometric function. The values are tabulated and used in treatment planning.

Up to now, the variation of the attenuation and scatter has not been modulated according to the angular angle θ of the point of interest. In the TG43 formulation this is taken into account by anisotropy function F(r, θ). For a perfect point source, the values of the radial dose function could be used for all angles θ; in other words, it would be perfectly isotropic or F(r, θ) = 1 for all r and θ. However, because of the linear nature of clinical sources and variation in the composition of internal components and encapsulation (which can modify the energy spectrum), including the drive wire if applicable, the radial dose function must be modulated.

Eq 27

In this equation, the angular dependence dose rate at a given distance r is given relative to that at 90° (the definition for g_X[r]), corrected by the ratio of the geometric factors when going from θ₀ to θ. It should be obvious to the reader that the anisotropy function is always 1 for all r when θ = θ₀. In PSIs, it is not practical in the current state of technology to consider the individual orientation of each seed for the final dose calculation. First, the seeds are small and these orientations are difficult to extract from CT or MRI. Second, it would have to be done consistently for anywhere from 50 to more than 120 seeds depending on the prostate volume and seed activity used. In such cases, correction of dose rate of each seed from 2-D anisotropy function is not relevant. It is, however, possible to average the anisotropy at a given r over 4π (the complete sphere). This is the 1-D approximation of the anisotropy function, which becomes ψ(r).

To conclude this section, let us summarize by providing the possible revised TG43 equations:

Eq 28

Eq 29

Eq 30

Equations 28 and 29 are recommended. Note that the first 1-D formulation uses the line source geometric factors and radial dose function. This is the recommended 1-D formulation by the AAPM TG43 because it provides the greatest accuracy at small distances for sources that are not perfect point sources. It is, however, not systematically implemented in commercial planning systems. A source can be considered as a point source for distances more than three times the active length of the source (r > 3L). The second 1-D formulation uses the point source approximation for the geometric factor. This is the 1-D formulation used in most planning systems at this time. Historically, there was also a 1-D formulation that averaged the anisotropy factor not only over all angles, but also over all distances to obtain the anisotropy constant ψ_an.

Clinical Considerations

The TG43 formalism was originally intended for low-dose-rate (LDR), low-energy seeds used in PSIs. However, as we have shown previously, the dose rate equation has parameters that depend specifically on the source design. As such, it is now widely used for HDR and PDR brachytherapy.

In 1998, an Ad Hoc Subcommittee of the Radiation Therapy Committee of the AAPM published recommendations on the prerequisite for clinical use of new low-energy photon sources. These recommendations advocate that:

1 It should be the manufacturer’s responsibility to provide the end user with air-kerma strength calibrated sources. This calibration should be traceable to National Institute of Standards and Technology (NIST) primary air-kerma strength standard. Furthermore, it is the vendor’s responsibility to ensure periodic validation of its source strength by an independent (primary or accredited) laboratory.

2 The dosimetric characteristics of a new source should be determined by at least one, preferably two, independent experimental studies. In addition, a detailed Monte Carlo study of the source should be conducted. Both the experiment and Monte Carlo studies taken together should enable all relevant dosimetric parameters of the TG43 formalism to be extracted. These recommendations also state that these studies should be published in a peer-reviewed journal.

3 Radiological Physics Center (RPC) maintains a registry of brachytherapy sources that meets the AAPM dosimetric prerequisites.¹⁰ Relevant literature for each source, along with a description of the AAPM prerequisites, are also given on the registry website.

This systematic process has not been officially implemented for HDR and PDR sources at this time. As such, the source registry is limited to iodine and palladium sources currently. However, the practice of performing experimental measurements and Monte Carlo simulations on brachytherapy sources of any type can almost be considered a standard procedure in the scientific community.

Beyond these recommendations, it is the responsibility of the medical physicist to review periodically the scientific literature for published consensus or new dosimetric parameters for the sources used in clinics and to ensure proper entry of the parameters in the planning system. Finally, it is also the responsibility of the medical physicist to experimentally verify, using a calibrated well chamber (traceable to an accredited dosimetry calibration laboratory), the air-kerma strength against the vendor’s calibration. A recent report by the AAPM low-energy brachytherapy source calibration working group has made detailed recommendation in that regards.

Modern Brachytherapy Photon Sources

A number of isotopes have been used throughout the years for cancer treatments. Examples are given in Table 13-1. For radiation protection (safety) reasons, a number of nuclides are not used anymore, such as radium, which has a very long half-life and emits high-energy photons. Furthermore, high-energy sources have a large mean path length. Although this generates relatively uniform dose in the target volume, the penetrating photons are a disadvantage for the protection of OARs compared with the low-energy sources, such as iodine 125 (¹²⁵I) for example.

Table 13-1 Properties of Some Common Radionuclides Used in Brachytherapy

Thus, to produce sources that are easier to handle, to shield, and to subsequently dispose of, the industry has moved toward lower-energy and shorter-half-life isotopes. All the nuclides shown in the bottom portion of Table 13-1 are good examples of isotopes used in modern brachytherapy. For permanent prostate interstitial brachytherapy, the patient body itself constitutes enough shielding, in the energy range of 20 to 30 keV, that the patient is not a danger to his family (even though the “as-low-as-reasonably-achievable [ALARA] principle recommends not to hold small children directly on one’s lap for an extended period). The half-value layer of lead needed for ¹²⁵I photons is 0.025 mm. The latest incarnation of this trend is unequivocally the electronic sources. The latter are miniature x-ray generators, which produce no radiation when the current is off and are sufficiently low-energy to be used in a regular radiology suite.

In the following, the development of brachytherapy sources for PSIs and temporary implants is discussed further.

Permanent Seed Implants

In PSIs, multiple sources or seeds are used. The exact number and their positions within the treatment volume are the results of the planning process. It depends on the prescription dose, the seed strength, and the treatment volume. In this type of treatment, the seeds are left permanently, as the name describes, within the human body. Thus the radioisotopes are encapsulated in a biocompatible shell. Three radioisotopes are used clinically in PSI: ¹²⁵I, palladium 103 (¹⁰³Pd) and more recently cesium 131 (¹³¹Cs). According to the brachytherapy source register at the RPC,¹⁰ 16 seed models meet the general AAPM dosimetric prerequisites at this time. For most seed models, the shells or capsules are made of titanium (Ti), a well-known material in the medical field. All clinically used seeds are between 4.5 and 5 mm in physical length and 0.8 mm in diameter. The shell is usually laser-welded at the tips of the capsules. This can have an affect on the anisotropy function of the seed resulting from the extra attenuation of photon exiting through the welding. Encapsulation provides protection to body tissues or fluids, which are never in direct contact with the radioisotopes. Therefore, the seeds can be used in patients that have allergies to iodine, for example. Furthermore, the Ti melting point is 1941 K. Therefore, they are highly resistant to all types of sterilization procedures and to cremation as well, should a patient die within a few months after the implants are placed.

Recently, another type of shell made of biocompatible polymer was proposed. Polymer has the advantage of being composed of low Z elements (water-like material), attenuating the emitted photons. Consequently, less internal activity should be needed to produce a specific dose at the reference point. Another advantage is that interseed attenuation (ISA)—the attenuation of the photons from one seed passing through another seed—was also shown to be less.

The internal makeup of a brachytherapy seed varies between manufacturers. The radioisotopes are usually deposited in the form of silver halide on metal holders such as beads and rod, as an organic matrix or coated (ion exchanged or other processes) on resin or polymer beads.

A third important component is a high Z radiopaque marker for postinsertion radiologic identification. These can be in the form of rod, beads, and cylindrical bands. Silver, gold, gold-copper alloy, tungsten, and platinum-iridium alloy are the most widely used materials.

Three very different designs are portrayed in Fig. 13-1. The first (left) is a ¹²⁵I seed. The radioactive material is deposited on a silver rod (red) in a welded Ti capsule (black). This design is typical of the 6711 seed and of the SelectSeed. The total seed length is 4.5 mm for both models. The middle seed is model 2335 ¹⁰³Pd seed. The radioactive material is deposited on three polymer beads (red) on each side of a 1.2-mm long tungsten marker (yellow). A welded, double-wall Ti capsule is used (black) to seal the components. The total seed length is 5 mm (4.25 mm active length). Finally, the last design represented is the IBt OptiSeed 1031L ¹⁰³Pd seed. The innovative part of this seed is its biocompatible polymer shell made only of carbon and hydrogen (orange). A 2-mm long gold marker (yellow) is placed in the middle of the 5-mm long seed. On each side of the marker, a palladium tube (red) 0.7 mm in length emits the low-energy photons. The active length is 3.7 mm.

FIGURE 13-1 • A Monte Carlo study of seed design on the interseed attenuation in permanent prostate implants.

(From Afsharpour H, D’Amours M, Coté B, et al: A Monte Carlo study on the effect of seed design on the interseed attenuation in permanent prostate implants, Med Phys 35:3671, 2008.)

Let us stress that many more designs exist.¹¹ Because of the internal configuration of the marker and nuclides, and the use of different marker and encapsulation materials, it should be obvious to the reader that each seed will have specific photon emission patterns: intensity as a function, polar angles, and energy spectrum. Therefore, we emphasize once more, that TG43 dosimetric parameters are different for each seed.

Iodine 125

¹²⁵I used in an LDR source for brachytherapy is produced mainly in a neutron capture process, through xenon 124 (¹²⁴Xe) gas target: ¹²⁴Xe(n, γ)¹²⁵Xe. In turn, ¹²⁵Xe decays into ¹²⁵I by electron-capture (EC) transition: ¹²⁵Xe → ¹²⁵I + ν + E_b. The half-life of the ¹²⁵Xe is 16.9 hours. ¹²⁵I also decays by EC into an excited state ¹²⁵Te*, producing the maximum photon energy of 35.5 keV by gamma decay (6.7% of the time). In addition, the transition leads to characteristic x-rays of energy between 27.2 to 31.7 keV (K-shells) as a result of internal conversion (93.3%). Note that a very low-energy x-ray of 3.8 keV is also possible (15%), but in practice is attenuated within the source capsule. These low-energy x-rays are omitted in the following. The same is true for the various electrons emitted in the decay process; these are low-energy and are filtered out by the various components of the source.

The half-life and photon energies are given in Table 13-1. The specific activity of ¹²⁵I is more than 600 GBq/mg⁴ and iodine seeds are currently being used not only for PSIs (with air-kerma strength between 0.4 U to 0.8 U) but also eye plaque (up to 5 U).

Palladium 103

¹⁰³Pd can be produced in the neutron activation of ¹⁰²Pd: ¹⁰²Pd(n,γ)¹⁰³Pd and by nuclear reaction with a proton beam: rhodium 103 (¹⁰³Rh)(p,n)¹⁰³Pd. Because stable ¹⁰²Pd occurs only at 0.9% level, an enrichment process (and potentially a further separation process following the activation) is needed to obtain high specific activity of ¹⁰³Pd via neutron activation. The other avenue, p-n nuclear reactions, has the advantage of using ¹⁰³Rh (natural abundance 100%). In practice, this isotope can be produced with a very high specific activity, more than 2500 GBq/mg.

¹⁰³Pd decays by EC to excited states of ¹⁰³Rh, 99.9% to the first excited state. Multiple gamma decays are possible from the excited states (39.7 keV to 497 keV). However, the emission probabilities for the gamma are all less than 0.07%. Internal conversion is the preferred mode of de-excitation from the first excited state to the ground stage. The probable emission energies for this isotope are the resulting characteristic x-ray emission between 20 to 23 keV. Overall, the mean photon energy of a ¹⁰³Pd seed is lower than ¹²⁵I (21 versus 28 keV). However, its shorter half-life results in higher initial dose rate compared with ¹²⁵I. ¹⁰³Pd has been promoted to be more biologically effective against rapidly proliferating tumors; however, limited support can be found for this notion. Up to now, clinical results do not show significant difference in survival rate.

Cesium 131

Two modes of production have been used to obtain ¹³¹Cs radioisotopes. Both first produced the parent nuclide barium 131 (¹³¹Ba), which decays into ¹³¹Cs by EC.

1 Neutron activation: ¹³⁰Ba(n,γ)¹³¹Ba

2 Nuclear reaction: ¹³³Cs(p,3n)¹³¹Ba ¹²

The natural abundance of ¹³⁰Ba is small (0.11%) and ¹³²Ba (0.10%) will lead to ¹³³Ba, a long-lived radioisotope (10.52 years) that decays by EC to ¹³³Cs with a high proportion of gamma rays between 53 and 384 keV. Note that ¹³⁴Ba, ¹³⁵Ba, ¹³⁶Ba, and ¹³⁷Ba are all stable isotopes. Thus, ¹³⁰Ba target enrichment is necessary. On the other hand, ¹³³Cs is a stable isotope with 100% natural abundance. Therefore, the nuclear reaction (p,3n) appears to be an interesting alternative. However, ¹³³Cs(p,n)¹³³Ba reaction is also possible. However, the production cross-section for (p,3n) reaction can be significantly higher than for (p,n) by choosing the appropriate incident proton beam energy.

Independent of the production mode, ¹³¹Ba decays into the daughter product ¹³¹Cs by EC ( = 11.50 days). ¹³¹Cs also proceeds by EC in ¹³¹Xe with a half-life of 9.69 days. This last process results in low-energy x-rays between 29.46 keV and 34.42 keV. There is no gamma emission.

Note that this isotope possesses the shortest half-life of the three isotopes available for PSI. Thus the initial dose rate for ¹³¹Cs is higher than that of ¹⁰³Pd and ¹²⁵I.¹³^,¹⁴ It is thus more susceptible to postimplant edema. Finally, the controversies remain in comparing the dose to both the prostate and the OARs with different radionuclides, in particular when the biologic effective dose is considered.¹⁵

High-Dose-Rate and Pulsed-Dose-Rate Brachytherapy

HDR brachytherapy treatments deliver doses of more than 12 Gy/hr. As such, HDR brachytherapy treatment usually lasts a few minutes. For its obvious advantages, our interest is toward afterloading delivery. Afterloader units are composed of an encapsulated radioactive source (Ti or stainless steal) at the tip of a stainless steel drive wire. While not in use, the source sits into a shielded area of the unit. The unit is controlled from outside of a shielded treatment room, decreasing direct manipulation and increasing safety. For treatment, the source travels through various types of applicators or inside needles and catheters. The source can stop at various positions, called dwell positions, for a given amount of time, called dwell time. The determination of the individual dwell position and dwell time is part of the planning process. Physicians have used modulations of these degrees of freedom for decades before the advent of external-beam, intensity-modulated radiation therapy. However, the full extent can only be explored with the recently available inverse planning algorithms.

A special case of HDR is PDR brachytherapy. In PDR, the treatment time is limited to brief irradiation for a given period to mimic LDR brachytherapy, in the range of 1 to 3 Gy/hr. The same type of equipment is used.

The most widely used nuclide for HDR or PDR brachytherapy is iridium 192 (¹⁹²Ir). Compared with cobalt 60 (⁶⁰Co) or ¹³⁷Cs, ¹⁹²Ir’s lower average photon energy makes it easier to shield. This offsets the disadvantage of a shorter half-life. Furthermore, the high specific activity of ¹⁹²Ir (341 GBq/mg) compared with ¹³⁷Cs allows for its use for both LDR and HDR brachytherapy with relatively small sources.

Iridium 192

¹⁹²Ir is produced in the nuclear reactor in the reaction ¹⁹¹Ir(n,γ)¹⁹²Ir. ¹⁹¹Ir composes 37.3% of natural iridium, ¹⁹³Ir making 62.7%. The material can be enriched so as to reduce the contamination of radioactive ¹⁹⁴Ir from ¹⁹³Ir and increase ¹⁹²Ir production. Interestingly, ¹⁹²Ir has a high neutron capture cross section, allowing double-neutron capture by ¹⁹¹Ir into ¹⁹³Ir. Hence, the production rate of ¹⁹²Ir is limited by its own decay and by its transformation into ¹⁹³Ir.

¹⁹²Ir decays into platinum 192 (¹⁹²Pt) via β⁻ decay 95.1% of the time and the remaining 4.9% in osmium 192 (¹⁹²Os) by EC. This leads to a complex decay pattern resulting in 29 gamma emission peaks from 0.110 to 1.378 MeV, various characteristic x-rays, and numerous electrons up to 1.377 MeV. For the β⁻ decay channels, de-excitation of ¹⁹²Pt* to ¹⁹²Pt (ground state) proceeds mainly via gamma emission and in a smaller fraction by internal conversion. Gamma rays of 0.296, 0.309, 0.317, and 0.468 MeV are the most probable (29% to 83%). This decay chain leads to 2.2 photons per β⁻ decay. The second, less probable, branch leading to ¹⁹²Os* by EC also decays mainly via gamma emission to the ground state. The most probable energies are the 0.206 and 0.485 MeV, but they make up only a small percentage.

The active ¹⁹²Ir core is usually made of an Ir-Pt alloy (10%-30% Ir and 70%-90% Pt) to make it more mechanically robust. The core itself is approximately 3 to 4 mm long and 0.3 to 0.4 mm in radius.

Ytterbium 169

Battista et al.¹⁶ have proposed ytterbium 169 (¹⁶⁹Yb) as an alternative radionuclide for brachytherapy. Its specific activity is approximately 2.5 times higher than ¹⁹²Ir, making it a good candidate for production of a small source of high strength. With an average energy of 93 keV, ¹⁶⁹Yb would lower the shielding requirements of an HDR brachytherapy room. This nuclide is produced in a reactor from the reaction ¹⁶⁸Yb(n,γ)¹⁶⁹Yb. Because ¹⁶⁹Yb composes only 0.13% of natural Yb element, an enrichment process must be performed. ¹⁶⁹Yb decays by EC in excited states thulium 169 (¹⁶⁹Tm). Return to the ground state proceeds via gamma emissions (45%) or internal conversion (55%). Seventy-four γ-ray transitions (from 63 keV to 782 keV), 13 characteristic x-rays, and more than 100 different electron energies (the most probable all being less than 130 keV) are emitted. The most common γ-rays are 63.12 keV (44%), 109.8 keV (17.5%), 130.5 keV (11.3%), 177.2 keV (22.2%), 198 keV (35.8%), and 307.7 keV (10%). A potential downside of this isotope from a clinical perspective is that its half-life is less then 50% of that of ¹⁹²Ir, and it therefore needs more frequent source change.

Thulium 170

Thulium 170 (¹⁷⁰Tm) is an interesting radioisotope, first because of its half-life of 128.6 days seems practical from a clinical perspective. Second, the average x-ray energy of 66 keV, lower than ¹⁶⁹Yb and ¹⁹²Ir, would allow for easier shielding and greater protection for the OARs. Finally, its specific activity is only 30% lower than ¹⁹²Ir. This enables the making of compact sources with high strength.

This radioisotope is made by neutron activation through the reaction ¹⁶⁹Tm(n,γ)¹⁷⁰Tm. The abundance of stable ¹⁶⁹Tm is 100%. ¹⁶⁹Tm is the rarest of the rare-earth elements (although found in quantities similar to silver) composing less than 0.01% of the ores from which it is extracted.¹⁷ Thus obtaining high purity materials is expensive. The decay of ¹⁷⁰Tm proceeds mainly (99.87%) via β⁻ decay. It can decay directly to the ground state (81.6% of the time) with a maximum beta energy of 968 keV (average energy of 323 keV). 18.3% of time the β⁻ decay will go to the first excited state and the associated beta energy will be lower (maximum of 883 keV, average of 290.5 keV). Final decay to the ground state can proceed directly by gamma emission of 84.25 keV photons, 2.48% of the time or via internal conversion leading to a series of six x-rays and six low-energy electrons. Apart from the 84.25 keV, the total x-ray emission probability is small at 6.25% for energies ranging from 7.42 keV (2.93%) to 60.96 keV (0.12%).

The other decay branch is EC to erbium 170 (¹⁷⁰Er). The probability of this decay is low at 0.13%. The x-rays (seven rays between 6.95 and 78.7 keV) and low-energy electron (six from 5.5 to 78.3 keV) spectra are a factor of 10 to 100 less intense than those from the main decay branch.

However, it should be noted that because of the low gamma and x-ray emission intensities, the air-kerma rate constant for ¹⁷⁰Tm is approximately 200 times lower than ¹⁹²Ir.⁹ Thus, a thulium source might very well be limited to PDR applications. Moreover, the effect of the high-energy electron emitted in beta decay probably needs to be taken into account both in terms of bremsstrahlung emission¹⁸ and also in terms of the possibility of insufficient encapsulation thickness to stop all betas. In such a case, an increased dose could be expected in the first few mm outside of the source capsule.

Electronic Sources

The latest in the field of brachytherapy are electronic sources.¹⁹ These are composed of a miniature x-ray source, which emits x-ray only when the tube current is on. Xoft has commercialized the first electronic sources, Axxent, and these are Food and Drug Administration–approved for specific treatments. It is a low-energy source of 50 kV. As such, the dose fall-off is greater with electronic sources than ¹⁹²Ir. Because the energy is lower than current radiology type devices, the body itself constitutes sufficient shielding. This allows for treatment personnel to be present in the room with the patient during treatment. So in principle no special (shielded) treatment room is needed. Finally, there is no radioactive waste handling.

As with any x-ray device, an electronic brachytherapy source produces a continuous energy spectrum. For the Axxent model, that spectrum has a mean energy at or lower than that of ¹⁰³Pd,²⁰ but a dose rate high enough for HDR treatments. The x-ray tube is 2.25 mm in diameter and the complete assembly is 5.4 mm in diameter. The tube is water-cooled and disposable after usage.