Title Page
 Table of Contents
 List of Tables
 List of Figures
 Background and related researc...
 The design and construction of...
 Description and results of tests...
 Determination of the effect of...
 Conclusion and discussion
 Biographical sketch

Title: design and construction of a miniature dosimeter for the study of the effects of air cavities in radiation therapy
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Permanent Link: http://ufdc.ufl.edu/UF00090218/00001
 Material Information
Title: design and construction of a miniature dosimeter for the study of the effects of air cavities in radiation therapy
Series Title: design and construction of a miniature dosimeter for the study of the effects of air cavities in radiation therapy
Physical Description: Book
Creator: McKetty, Marlene Hope Patricia.
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Bibliographic ID: UF00090218
Volume ID: VID00001
Source Institution: University of Florida
Holding Location: University of Florida
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Resource Identifier: alephbibnum - 000074788
oclc - 04718089

Table of Contents
    Title Page
        Page i
        Page ii
    Table of Contents
        Page iii
        Page iv
    List of Tables
        Page v
    List of Figures
        Page vi
        Page vii
        Page viii
        Page ix
        Page x
        Page 1
        Page 2
        Page 3
    Background and related research
        Page 4
        Page 5
        Page 6
        Page 7
        Page 8
        Page 9
        Page 10
        Page 11
        Page 12
        Page 13
        Page 14
        Page 15
        Page 16
        Page 17
        Page 18
        Page 19
        Page 20
        Page 21
        Page 22
        Page 23
        Page 24
        Page 25
        Page 26
        Page 27
        Page 28
        Page 29
        Page 30
        Page 31
        Page 32
    The design and construction of the dosimeter and peripheral apparatus
        Page 33
        Page 34
        Page 35
        Page 36
        Page 37
        Page 38
        Page 39
        Page 40
        Page 41
        Page 42
        Page 43
        Page 44
        Page 45
        Page 46
        Page 47
        Page 48
        Page 49
        Page 50
        Page 51
        Page 52
        Page 53
        Page 54
        Page 55
        Page 56
        Page 57
        Page 58
        Page 59
        Page 60
        Page 61
        Page 62
    Description and results of tests performed on the dosimeters
        Page 63
        Page 64
        Page 65
        Page 66
        Page 67
        Page 68
        Page 69
        Page 70
        Page 71
        Page 72
        Page 73
        Page 74
        Page 75
        Page 76
        Page 77
        Page 78
    Determination of the effect of air cavities
        Page 79
        Page 80
        Page 81
        Page 82
        Page 83
        Page 84
        Page 85
        Page 86
        Page 87
        Page 88
        Page 89
        Page 90
        Page 91
        Page 92
        Page 93
        Page 94
        Page 95
        Page 96
        Page 97
        Page 98
        Page 99
        Page 100
        Page 101
        Page 102
        Page 103
        Page 104
        Page 105
        Page 106
        Page 107
        Page 108
        Page 109
        Page 110
        Page 111
        Page 112
        Page 113
    Conclusion and discussion
        Page 114
        Page 115
        Page 116
        Page 117
        Page 118
        Page 119
        Page 120
        Page 121
        Page 122
        Page 123
        Page 124
        Page 125
        Page 126
        Page 127
    Biographical sketch
        Page 128
        Page 129
        Page 130
Full Text


By *






The author wishes to acknowledge with gratitude the help and

guidance given by her chairman and the members of her supervisory

committee in her research and writing of this manuscript. She would

also like to thank Dr. Million of the department of Radiation Therapy

and Dr. Carroll of the department of Nuclear Engineering Sciences for

providing funds to procure some of the equipment used in this research.

The author also wishes to thank the National Fellowships Fund for

assistance in defraying the cost of the typing of this manuscript.

Finally, thanks to Mrs. Janet Eldred for typing the manuscript, to

Mr. Wesley Bolch for his help with the graphics and to Dr. Robert

Luthman for his help with the photography.




LIST OF TABLES . . . . . . . . . . . .

LIST OF FIGURES . . . . . . . . . . . .

ABSTRACT . . . . . . . . . . . . . .


1 INTRODUCTION . . . . . . . .


Types of Dosimeters . . . . . .

Ionization Chambers . . . . .
Thermoluminescent Dosimeters . . .

Solid State Dosimetry . . . . . .

PN Junctions . . . . . . . .
Use of PN Junctions . . . . . . .
Fabrication of PN Junctions . . . . .
P-I-N Detectors . . . . . . . .
Fabrication of P-I-N Diodes . . . . .
Methods of Detection and Measurement .
Uses of P-I-N Diodes . . . . . .
Problems Encountered in the Use of P-I-N
and PN Junctions . . . . . . .

Air Cavities . . . . . . .. . .
Objectives of Research . . . . . . .


Description of Ideal Dosimeter . . . . .
Prototype I . . . . . . . . .
Prototype II . . . . . . . . . .
Peripheral Apparatus . . . . . . . .

. . 9

. . 10
. . 12
. . 15
. . 16
. . 20
. . 21
S. 26

. . 27



S. 4





Amplifier Circuit ... ... . . . . . . .. 50

Operational Amplifier . . . . . . ... 50
Feedback Capacitor . . . . . . ... 53
Reset System . . . . . . . . ... .57
Circuit Construction . . . . . . ... 58

Display . . . . .. . . . . . . 59

THE DOSIMETERS . ........ . . . . ... 63

Tests on the Dosimeters . . . . . . .. 63

Reproducibility . . . . ...... ... ... 63
Linearity . . . . . . . . ... . 63
Temperature Dependence . . . . . ... 63
Directional Dependence.. . . . . . 71
Construction Parameters Affecting
Response . . . . . . . .... . 74
Energy Dependence . . . . . . . ... 75
Dose Rate Dependence . . . .... ... 76

Tests on Peripheral Equipment . . . . . ... 77

Cable Effect . . . . . . . .... .77
Position of Amplifier . . . . . ... . 78
Gain Settings . .. . . . . . . .... 78


Depth Dose Measurements . . . . . ... 80
Backscatter Measurements . . . .. ... . 93
Effect of Width of Air Cavity . . . . ... 103
Measurements in Anatomical Phantom . . .. 111

6 CONCLUSION AND DISCUSSION . . ... . . ... . 114

Physical Factors . . . . . . . ... 114
Use of Dosimeter. .. ... . .. .. . . 115
Air Cavity Measurements . . . . . ... 116

REFERENCES . . . . . . . . . . . . . 121

BIOGRAPHICAL SKETCH .. .. . . . . . . . . .... 128



1 Summary of Measurements with Commercial Diodes .

2 Gain Characteristics of the Amplifier System . .

3 Correction Factors for Increased Response of
Dosimeter 2 to Temperature . . . . ... .

4 Directional Sensitivity Measurements for
Dosimeter 2 . . . . . . . . . .

5 Dose Rate Dependence of Dosimeter 2 . . . .

6 Comparison of Silicon Diode and Ionization Chamber
Central Axis Depth Dose Measurements . . . .

7 Comparison of Change in Percentage Depth Dose
Caused by Air Cavity in a 6 x 6 cm2 6Co Field .

8 Comparison of Change in Percentage Depth Dose
Caused by Air Cavity in a 6 x 6 cm2 17 MeV
Photon Field . . . . . . . . . .

9 Results of Measurements in an Anatomical
Phantom . . . . . . . . . . .


. . 34

. . 56

. . 71

. . 74

. 77

. . 94

. . 110

. . 111

. . 112


Figure Page

1 Diagram of a p-i-n Chip . . . . . . ... 18

2 Diagram of Idealized Detector . . . . .... 37

3 Diagram to Illustrate Arrangement of p-i-n Chips in
Prototype 1 .. .. . . . . . ..... ... . 39

4 Photograph of Prototype 1 . . . . . . ... 42

5 Photograph of Prototype 2 . . . . . .... 46

6 Diagram of Requirements for Displaying Signal ..... 47

7 Typical Integrating Circuit . . . . . . . 49

8 Operational Amplifier Integrating Circuitry . . .. 52

9 Photograph of Amplifier and Display System . . . 62

10 Graph to Demonstrate Linear Response of Dosimeter 1
to Radiation . .. . . . . . . . . . 65

11 Graph to Demonstrate Linear Response of Dosimeter 2
to Radiation . .. . . . . . .... ... . 67

12 Temperature Dependence of Dosimeter 2 . . . ... 70

13 Directional Sensitivity of Dosimeters 1 and 2 .... 73

14 Depth Dose Measurements for 6Co ... . . . 82

15 Depth Dose Measurements for 60Co . . . . .. . 84

16 Depth Dose Measurements for 6Co .. . . . .. . 86

17 Depth Dose Measurements for 8 MeV Photons . . ... 88

18 Depth Dose Measurements for 8 MeV Photons . .... .. 90

19 Depth Dose Measurements for 17 MeV Photons . .... . 92



20a Diagram of Phantom Used for Backscatter
Measurements . . . . . .. .

20b Diagram of Phantom Used for Backscatter
Measurements . . . . . . .

20c Diagram of Phantom Used for Backscatter
Measurements . . . . . . .

21 Results of Backscatter Measurements .

22 Results of Backscatter Measurements .

23a Diagram of.Phantom Used for Air Cavity
Measurements . . . . . . .

23b Diagram of Phantom Used for Air Cavity
Measurements . . . . . . .

24 Results of Air Cavity Measurements . .

25 Summary of Air Cavity Measurements . .


S .. .. 96

S. . . 97

S. . . 98

. . . 100

S. . . 102

S. . . 105

. . 107

. . . . 109

. . . 119

Abstract of Dissertation Presented to the Graduate Council
of the University of Florida in Partial Fulfillment of the Requirements
for the Degree of Doctor of Philosophy



Marlene Hope Patricia McKetty

August 1978

Chairman: Lawrence T. Fitzgerald
Major Department: Nuclear Engineering Sciences

A miniature dosimeter utilizing silicon p-i-n chips was designed

for performing in vivo measurements in radiation therapy. The dosimeter,

which is less than 2 mm in diameter, can be inserted into patients and

phantoms for rapid measurements. An important advantage of this type of

dosimeter over thermoluminescent dosimeters and miniature ionization

chambers is the instant readout and the elimination of a waiting period

between measurements.

Dosimeters which have pn junctions are useful for detecting alpha

particles and fission fragments but are usually unsuitable for photons

because insufficient energy is deposited in the depletion region to pro-

duce an adequate signal. By using p-i-n diodes rather than diodes with

pn junctions, electrons and photons can be .detected since sufficient

energy can be deposited in the intrinsic region to produce a measurable


Different arrangements of the p-i-n chips were explored in an

effort to obtain minimal detector size, directional independence and a

large signal. The physical parameters of the dosimeters were examined.


These include temperature effect, directional sensitivity, effect of

encapsulating material, energy and dose rate dependence and linearity

of response.

For initial measurements in the study, the dosimeter, which was

operated in the photovoltaic mode, was used with a Keithley 616 digital

electrometer. Once the magnitude of the signal was established, an

electrometer was designed and built specifically for use with the diode

dosimeters. The electrometer allowed a voltage proportional to the

dose received by the dosimeter to be integrated and then displayed on

a digital panelmeter. A voltage divider network incorporated in the

integrating circuit was used to vary the effective feedback capacitance

and, therefore, the sensitivity of the system.

The dosimeters, which were calibrated against a National Bureau of

Standards calibrated Victoreen ionization chamber, were used to investi-

gate some of the effects of air cavities in a radiation treatment,

volume. These include change in backscatter and build-up characteristics.

Geometrical as well as anatomical phantoms of tissue equivalent material

were used for measurements. Cobalt-60 gamma rays and 8 and 17 MeV

photons from a linear accelerator were used as the radiation source for

the various measurements. Measurements were taken to determine the

variation in dose caused by changing the dimensions and geometry of an

air cavity.

In patient treatment planning, the presence of an air cavity in a

treatment volume is not considered when dose distributions are

determined. The measurements with the miniature dosimeter indicated

that an air cavity had a greater impact if a volume were being irradi-
ated with 60Cobalt rather than 17 MeV photons. As the dimensions of the
ated with Cobalt rather than 17 MeV photons. As the dimensions of the

treatment field and the length of the air cavity were increased, the

discrepancy between measured dose and calculated dose became more marked.

Since air cavities are often in the treatment volume of head and neck

cancers, these measurements suggest that the use of high energy photons

may be advantageous in obtaining a more uniform dose distribution.


In the treatment of cancer patients with radiation, a major problem

is the accurate determination of the dose at different points in the

treatment volume. The objective of treatment planning in radical radio-

therapy is to design a plan in which a tumorcidal dose is given to the

malignant tumor without surpassing the tolerance of the normal tissue.

The greater the dose received by the tumor, the greater is the chance

of destroying the malignant cells. A complete treatment plan should

provide an accurate representation of the doses that will be received

in an irradiated area. The methods of collecting the depth dose data

that are used for treatment plans are very precise; therefore, the basic

data with which one begins are accurate. However, the accuracy of the

depth dose data collected in a homogeneous medium, usually water, is

reduced when it is applied to a nonhomogeneous one.

It is difficult to ascertain the degree of accuracy needed in

radiotherapy. Direct experimental evidence relating degree of precision

and result of radiation therapy is unavailable. Therefore, one must

rely on data extrapolated from animal studies or examine clinical data.

With clinical data one must relate tumor control probability, or the

frequency with which normal tissue tolerance is exceeded, to the total

dose. These methods are inadequate; however, for lack of better methods

they are still in use.

Herring and Compton (1971) stated that the dose at the tumor or

other critical volume should be known to within 5 percent or greater

accuracy. This value was determined by the development of a mathe-

matical model which depicted the response of tumors to a course of

radiotherapy. The model illustrated that small variations in the radia-

tion dose had a large influence on the probability of cure. If the dose

is not known to this accuracy, there is an increased probability of

necrosis or lack of local control. Shukovsky (1970) demonstrated the

large effect of variations in dose of 5 percent or more on the local

control of tumors in the head and neck region. Fowler (1963) also

agrees that the accuracy for radiotherapy should be no worse than 5

percent but an accuracy of 2 percent is desirable. Dosimetric errors

should be small compared with individual clinical variations.

Overall, many aspects for obtaining accurate dose measurements

have been refined. However, there are still some aspects of measure-

ments which need considerable improvement. One aspect is the measure-

ment of doses at a specific point in a treatment volume with a dosimeter.

A high degree of accuracy is essential when measuring the change in dose

over a small area or for in vivo studies but there are few types of

dosimeters available for these measurements. There are often rapid

changes in the dose distribution depending on the anatomy of the irradi-

ated area and the type and energy of radiation. A well known example is

the build-up of dose from a Cobalt-60 (60Co) beam in soft tissue, the

dose changes from 30 percent of the maximum to 100 percent in the first

5 mm and then decreases slowly. Another area which requires improvement

is the correction of dose for the presence of internal and external

inhomogeneities. On entering the body, a radiation beam often does not'

pass through a solely homogeneous medium but may pass through air cavi-

ties as well as through different types of tissue (e.g., lung, bone, fat

and muscle), each of which affects the beam differently. The dose dis-

tribution is, therefore, different from what is shown on a treatment

plan where it is assumed that the beam is passing through a homogeneous

medium and that there is always sufficient scattering material.

A miniature dosimeter could be used for checking the doses at

different points in the treatment volume, especially in areas where

there is a known discrepancy between the dose received and the dose

shown on a treatment plan. This discrepancy is often manifested clini-

cally, e.g., a reaction may be more severe than would be expected for a

certain dose.


Types of Dosimeters

Ionization Chambers

The type of dosimeters most commonly used in radiation therapy are

ionization chambers which use the detection principle of gas ionization.

They are used for routine calibrations of therapy machines and for depth

dose and exposure measurements. Ionization chambers will continue to be

used because they can be made energy independent over a wide range and

they provide the most reliable data. Basically, an ionization chamber

consists of a small volume of gas and a pair of electrodes, one of which

is positive, the other negative. The wall of the chamber is lined with

material that serves as one electrode and the other electrode passes

down the middle of the chamber. When the air becomes ionized by radia-

tion the charges produced are collected at the electrodes. Under

proper conditions, e.g., at saturation voltage, the charge collected is

directly proportional to the amount of radiation hitting the ionization

chamber. The charge liberated in the chamber produces a voltage drop

between the electrodes and is measured with an electrometer. Errors

caused by attenuation of the radiation in the wall of the chamber may

be minimized by special design and calibration of the ion chamber. The

chamber wall and electrodes are of different chemical nature from the

air volume and, therefore, will contribute greater or lesser amounts of


photoelectrons than air. Thus, there is a dependence of calibration on

wall material and on the beam spectrum. This energy dependence is mini-

mized by making the wall relatively small and of a material similar to

air. Thewall is usually made of bakelite coated with graphite.

Victoreen R-meters are standard equipment in most radiotherapy

departments. Chambers of different sensitivities and with walls of

different thicknesses are used to measure low, medium and high energy

photons. According to Johns and Cunningham, "The Victoreen is capable

of great precision, but it is useless to measure exposure at a 'point'

because of its size and should certainly.not be used to measure expo-

sures in body cavities" (1969, p. 238). Apart from the fact that the

size of these dosimeters prohibits their insertion into body cavities,

they only measure the mean value of a radiation beam at any point of


The design of miniature ionization chambers to measure exposures

in body cavities has been described (Skoldborn, 1959). Several require-

ments were specified for the first such chambers to be made. The

radiation measurement requirements are

1. air equivalent material around the measuring volume,

2. energy independent sensitivity,

3. registration of radiation independent of direction of rays and

4. tissue equivalent material in the entire chamber in order not

to disturb distribution of rays.

The electrical requirements are

1. high insulating properties and tolerance to radiation of

insulating material used,

Victoreen R-meters are manufactured by the Victoreen Instrument
Company, Cleveland, Ohio.

2. lack of dielectric after-effect in insulating material and

3. sufficient capacitance to permit measurements within the

intended range at a suitable voltage.

The measuring requirements are

1. accuracy,

2. ease in handling of the chamber which should not be fragile


3. insensitivity of the chamber material to changes in temperature

between 15 and 40 C.

Most of these requirements were met in the design and construction of

Alderson Sievert ionization chambers which are supposed to approximate

point detectors. This type of chamber is 2 cm in length and 5 mm in

diameter with a measuring air volume of 25 mm3. It is designed for

measurements of doses ranging from one roentgen up to 300 roentgens.

Sievert chambers have been used for dose measurements in phantoms

and in patients. Although they have provided a useful service, Sievert

chambers are not really adequate for routine patient dosimetry. One

disadvantage is that they are not really small enough to be easily in-

serted in patients for in vivo dosimetry. Another disadvantage of these

chambers is that they should only be exposed a maximum of once daily.

After exposure, before the dose can be read on the electrometer, a wait-

ing period of about one hour is needed. The use of these chambers was

discussed briefly by Alderson et al. (1962) and more extensively by Dahl

and Vikterlof (1960). McKetty (1975) described the use of Sievert cham-

bers in a RANDO phantom to determine if the depth dose calculated at

certain points in a mantle field was the same as the measured dose.

These miniature ionization chambers are no longer manufactured in the


Thermoluminescent Dosimeters

Thermoluminescent dosimeters (TLDs) have been used for patient

dosimetry for the past three decades. Daniels (1950) proposed the fact

that thermoluminescence could be used as a method of dosimetry since

the amount of light emitted from certain materials on heating after

irradiation was proportional to the amount of radiation received.

Because of their small size TLDs could be used to approximate point

detectors and have been used for measuring doses in vivo and whenever a

small dosimeter is imperative. Cameron et al. (1968) have done exten-

sive research with lithium fluoride (LiF) dosimeters and this resulted

in the development of several types of LiF dosimeters, in particular

TLD 100 which is manufactured by Harshaw Chemical Company for routine

dosimetry. Lithium fluoride is also available as loose powder, as

powder sealed in tubes or as plastic (Teflon) sealed rods or discs

ranging in size from about 1 mm to several cm.

There are several advantages in using LiF as a dosimeter. These


1. the relative energy independence,

2. the good directional uniformity,

3.. minimal fading,

4. its small size and inert nature,

5. a response that is independent of dose rate and

6. the fact that the effective Z is close to that of tissue, thus

causing the response per roentgen to be similar to that of soft tissue.

Lithium fluoride shows a linear response with dose up to about 700 rads,

above this level the response is supralinear. The disadvantages of LiF

dosimeters are

1. the change in sensitivity with reuse,

2. the long preparation time needed before measurements,

3. the delay in getting a readout after exposure and

4. the destructive readout.

The accuracy of dose measurements with TLDs is at best 5 percent. They

lack the reproducibility needed to be accurate enough for exact dosage


Worsnop (1968) has described a study whose purpose was to deter-

mine whether the dose distribution that is actually delivered to

patients is the same as the dose prescribed. An Alderson RANDO phantom

containing LiF dosimeters distributed in a specific pattern was sent to

15 different institutions to be irradiated under certain conditions.

The institutions were participating in a national clinical study.

Hendee (1966) described a comparison of beta depth dose data obtained

with LiF dosimeters and by film and ionization chamber techniques.

Intracavitary measurements utilizing LiF have been discussed by several

researchers (Cameron et al., 1961; Naylor, 1967).

Despite the useful roles played by miniature ionization chambers

and LiF dosimeters in patient dosimetry, there are still many areas in

dosimetry which are beyond the scope of these detectors. Fowler (1963)

states that of the three main dose ranges requiring measurements it is

in the dose range used for radiotherapy of a few rads to several thou-

sand rads that small, reliable and convenient measuring devices are not

yet available. Several clinical applications in which other methods of

detection and dosimetry were used can now be performed with semicon-

ductor devices (Hertz and Gremmelmaier, 1960; Abson et al., 1968; Baily

and Norman, 1962). Solid state dosimetry may be able to solve some of

the problems that remain.

Solid State Dosimetry

Solid state dosimeters use the detection principle of electrical

conductivity and behave as solid state ionization chambers. A temporary

change in electrical conductivity occurs in semiconductor materials when

they absorb ionizing radiation--this change is a function of the

absorbed dose rate. Usually when semiconductor detectors are used in

medicine and biology, they are used as particle detectors and nuclear

spectrometers rather than as dosimeters. Marcus (1973) described the

use of silicon probes for physiological studies in the human body and

presented some of the reasons for their success and failure as radiation

detectors. Most procedures that use the probes for the detection of x

and y rays are unsuccessful, some success is reported, however, in the

detection of very low energy x rays as from plutonium localized in

lymph nodes and of g particles. Lauber and Wolgast (1972) also dis-

cussed the construction and properties of some miniature detector probes

for measurement of and y tracer activity in vivo.

The properties that are desirable for semiconductor detectors and

dosimeters have been discussed by several authors (Friedland and

Zatzick, 1967; Hendee, 1970; Fowler, 1963; Friedland and Katzenstein,

1973). The observed effect should be independent of the type of radia-

tion and the dose rate. It should vary linearly with the energy

deposited and be independent of temperature. The detector should have

a short time constant and variable sensitive depth. It should be small

enough to allow for convenient insertion into the body and have a

protective coating thick enough to safeguard against attack from body

fluid but thin enough to enhance sensitivity. It should be operated at

low potential to minimize danger to the patient.

Silicon and germanium are the two most important semiconductor

materials. They both have four electrons located in the valence band

at absolute zero temperature. When the temperature is increased some

of the valence electrons will acquire enough energy to cross the for-

bidden band or energy gap and enter the conduction band, thereby leaving

a hole in the valence band. The hole is considered as having a positive

charge. The free electrons and holes will act as current carriers if

there is an applied voltage. The magnitude of this current which is

called the drift current depends on the resistance of the semiconductor

material and the number of electron hole pairs that are thermally

generated. There is also a surface leakage current. The energy gap

for silicon is approximately 1.1 ev. Certain impurities are often

added to semiconductor material in a process called doping in order to

increase either the electron or hole conduction. Whereas an n type

impurity has five valence electrons and will result in a semiconductor

with excess electrons, a p type impurity has three valence electrons

and will lead to the formation of excess holes.

PN Junctions

The design of most semiconductor devices is based on the pn

junction, but there is also the p-i-n structure. A pn junction is

simply the transition from a p type material to an n type material

within a semiconductor crystal. Diffusion of majority carriers across

the junction causes a build-up of charge near the junction. This region,

which is called the depletion region or space charge region,has a posi-

tive charge on the n type side and a negative charge on the p type side.

In the absence of an applied voltage there is no net current across the

junction. However, if a voltage source is applied with the negative

terminal attached to the n type material and the positive terminal

attached to the p type material (forward bias) majority carriers will

move across the junction, i.e., holes will move from the p type side and

electrons from the n type side. If instead there is a reverse bias, the

majority carriers are pulled away from the junction and the depletion

area is widened. The only current that flows is due to minority

carriers. When the depletion region is subjected to ionizing radiation

new electron hole pairs are generated which can then be collected by

the potential across the semiconductor. These pairs form a current

which is similar to an ionization current in an ion chamber. The number

of charge carriers produced by radiation in the depletion region and

collected per unit time is proportional to the radiation energy absorbed

in the region and not to the number of absorbed photons. The following

equation is obtained

Nd = (AR gi/p) (1 exp (-pw)) (1)


Nd is the number of radiation produced charge carriers,
AR is the irradiated silicon surface area in cm ,
gi is the generation rate of electron hole pairs in cm3
sec ,

w is the width of the depletion region in cm and

p is the linear attenuation coefficient of the radiation
in silicon in cm .

The properties of a pn junction structure are determined by the

width of the depletion region and the carrier lifetime (T) in the base

material. The depth of the depletion region is proportional to

( (pV) where p is the base resistivity and V is the.applied bias

voltage, g = 1/2 for n type silicon and B = 1/3 for p type silicon.

The average energy (E) required to form an electron-hole pair in

silicon is 3.5 ev. This is approximately 1/10th of the amount of energy

needed to produce an ion pair in.an air-filled ionization chamber where

30-40 ev are needed. Brown (1961) pointed out that this low value for

C is attractive in a counting device because of the reduction in the

statistical fluctuations in the number of pairs produced. Because

silicon is approximately 1800 times as dense as air, the number of

electron pairs formed per unit volume is 18,000 times greater than in

an air-filled ionization chamber. These factors enable solid state

dosimeters to be of small volume. Baily et al. (1962) noted that the

advantages of using high density material for radiation detection had

led to extensive research on semiconductors, but the manufacture of

satisfactory detectors was a comparatively recent achievement. The

charge collection time in a semiconductor is much less than in an ioni-

zation chamber because of the greater charge carrier mobility and the

shorter distance for transit. Semiconductor dosimeters are more rugged

than ionization chambers and can be made into shapes to suit particular


Use of PN Junctions

PN junctions are used in diodes and in photovoltaic cells. Diodes

are normally used as rectifiers in electronic circuits but because of

their ability to produce a current when exposed to ionizing radiation

they can be utilized as dosimeters. Photovoltaic devices are used to

convert solar energy into electrical energy but can also be used as

dosimeters. Several authors have recommended the use of silicon detec-

tors for x and y ray dosimetry (Baily and Norman, 1963; Calkins, 1962;

Jones, 1963). The feasibility of using such dosimeters has been

substantiated (Scharf and Sparrow, 1966; Scharf, 1960; Scharf, 1967).

When silicon diodes and photovoltaic cells are used as dosimeters

several problems must be considered. Trump and Pinkerton (1967) list

some of these problems as

1. energy dependence of calibration in terms of absorbed dose,

2. effect of high atomic number material often used in mounting

the silicon wafer and

3. optimum electrical readout.

Other problems are the dependence of the signal on environmental

factors rather than solely on dose and angular and directional


Calkins (1962) discussed the fact that in low energy beams, energy

absorption as a function of beam energy was different in photovoltaic

cells and tissue. This is because of the difference in atomic number

of silicon and tissue. He indicated that for photons of energy greater

than 0.5 MeV, the photovoltaic cell reading could be converted into

absorbed dose by a constant multiplication factor. This is because

absorption is mainly by Compton interaction and, therefore, depends on

electron density per gm which is the same for silicon and tissue.

Trump and Pinkerton (1967) have reported the use of a commercially

available pn junction diode as a radiation probe for an automatic

isodosimeter. They measured dose distributions produced by high energy

radiotherapy sources, namely, 60Co y rays and 6, 10, 15 and 20 MeV

electrons. In 1962 Stanton and Lightfoot had reported that pn junction

diodes should be better suited for in vivo measurements than conven-

tional radiation detectors but that there was a lack of experience in

their use. They tested gallium arsenide diodes but concluded that

silicon diodes would show a smaller energy dependence because of their

lower atomic number. As early as 1958 Moody et al. reported the

response of silicon pn junctions to x and y rays. They determined that

a calibrated photovoltaic cell connected to a sensitive voltmeter by

shielded electrical leads were all that was required to build a y ray

dosimeter which would have a wide range of response. More experience

in the use of these dosimeters has been gained.

Raju (1966) reported the use of commercially available pn junction

diodes in determining beam profiles and depth dose distribution of a

cyclotron beam. Guldbrandsen and Madsen (1962) also experimented with

commercially available diodes and concluded that they fulfilled the

usual requirements of dosimeters, although dose rate measurements at

small energies were always more uncertain than those at higher energies.

Wright and Gager (1973) used a diode-scanner-recorder system to monitor

some operating parameters of a linear accelerator. They also used it

for measuring central axis depth dose data and off-axis data (Wright

and Gager, 1977). Because of problems that resulted from nonuniform

energy response of the diode, they developed an energy compensating

shield for the diode from a high Z material (Gager et al., 1977). The

shielded diode could then duplicate.the response of a Farmer probe.

Jones and Schumacher (1975) also developed an instrument utilizing

diodes to measure the output, energy and symmetry of a linear accelera-

tor beam. This device consists of a block of wood containing two

silicon diode radiation detectors along the central axis. The output

of each diode was amplified and displayed on a digital panel meter.

Parker and Johnson (1969) described a silicon pn junction detector

designed for checking the radiation dosage to the rectum and bladder of

patients with gynaecological cancer who are undergoing intracavitary

irradiation. Whelpton and Watson (1963) used pn junctions supplied as

silicon solar cells to measure the build-up curve from 60Co Y rays in

air. They found that the curve obtained using this method was similar

to that obtained using thin walled ionization chambers. Whelpton and

Watson (1963) designed a probe which was approximately quality indepen-

dent over the range 1.0 mm Cu HVT to 13 mm Cu HVT.

Fabrication of PN Junctions

PN junctions may be formed by two methods, surface barrier and

diffusion. Diffusion junction devices are formed by diffusing a suitable

impurity into the base material, that is, by diffusing a shallow layer

of a donor (or acceptor) material into high resistivity p type (or n

type) silicon. Goulding (1964) reports that phosphorus is usually used

as the diffusant and p type silicon as the bulk material. The diffused

layer is about 0.5 micron in thickness. The depletion layer forms in

the bulk p type silicon when a positive voltage is applied to the n

type layer. The depth of the depletion regions is given by the following


W = 0.32 (pV) micron (2)



W = depth of depletion layer,

p = resistivity of the p type material and

V = potential.

Ziemba et al. (1962) describes a conventional pn junction detector as

having four regions,

1. a thin phosphorus or lithium diffused layer,

2. a space charge region,

3. an undepleted silicon layer and

4. a boron diffused or aluminum alloyed layer which provides an

ohmic contact to the p type material.

Surface barrier detectors function in a similar manner to pn

junction detectors formed by diffusion but they are cheaper and easier

to construct (Parker and Morley, 1966) and more stable in their

properties. Because of their availability in a needle probe geometry,

they are well suited for many dosimetric applications in medicine.

Basically surface barrier detectors consist of high resistivity n type

silicon on one surface of which is a very thin p type layer formed by

spontaneous oxidation. Thus, a pn junction is formed close to the


P-I-N Detectors

Until fairly recently the most successful solid state radiation

detectors were surface barrier and pn junction silicon counters.

Whereas these detectors were adequate for alpha particles, low energy

protons, deuterons and fission fragments, they were unsuitable for use

with electrons, photons and minimum ionizing particles. This is

because insufficient energy was deposited in the shallow depletion

region to produce an adequate signal.

When semiconductor detectors are used for x and y rays the maximum

amount of energy should be deposited in the charge depletion region

(Baily and Mayer, 1961). That is, the photons should not be absorbed

until they have arrived at the depletion region and should be absorbed

before penetrating the depletion region. This may be achieved by

either increasing the atomic number of the semiconductor material or

increasing the depth of the depletion region. The first method is

achieved by using germanium (Z = 32) doped with gallium or zinc instead

of silicon (Glos, 1964). The latter method is used most often and can

be accomplished by either increasing the bias voltage or increasing the

resistivity of the semiconductor material. If the bias voltage is

increased the noise is increased and this is undesirable. One method

of increasing the resistivity is by diffusing lithium into silicon to

form a lithium drifted silicon detector which is also called a p-i-n

junction detector. Thus, the development of the p-i-n structure enables

solid state semiconductor detectors to be used successfully for x and y

ray and electron dosimetry.

A p-i-n junction consists of a wafer of extremely high resistivity

silicon on one surface of which is a thin p type diffusion and on the

opposing surface a thin n type diffusion (Figure 1). The high resis-

tivity material in the middle is the intrinsic (I) layer and it can

sustain a high electric field. The width of the intrinsic layer deter-

mines the electrical properties of the junction. The space charge

region will extend throughout the entire intrinsic region when a few

volts reverse bias is applied (approximately five volts) of if there is




Figure 1. Diagram of a p-i-n Chip

a high degree of compensation in the I region. The operating potential

can, therefore, be quite low.

If the space charge region extends throughout the I region there

is no variation of capacitance (C) with voltage, since the capacitance

is constant. At low voltage, however, the p-i-n detector behaves like

C a 1/V1/2 or C a 1/V1/3 (Ziemba et al., 1962). Since the capacitance

is quite low, p-i-n junctions can be used in applications requiring high

sensitivity. Mayer (1962) reported the results of an investigation to

determine the junction characteristics as a function of width of the

intrinsic region by the use of capacitance and current measurements.

The width (W) of the depletion region is given by

W = kA/4rC (3)


k = dielectric constant,

A = junction area and

C = junction capacitance.

P-I-N diodes may be characterized in several ways. The quantum

efficiency Nq refers to the number of electrons produced per photon.

The speed of response depends on what areas are irradiated, the amount

of reverse bias applied and the load resistance.

In order to detect both short and long wavelength photons, a diode

should have a thin p layer and a thick depletion area. This is because

long wavelength photons are absorbed near the surface while short wave-

length photons penetrate more deeply. The structure of p-i-n diodes

allows both long and short wavelengths to be detected in the same


Fabrication of P-I-N Diodes

Ziemba et al. (1962) described four methods for making p-i-n

detectors each of which has its advantages and disadvantages. These

methods are

1. paint on technique,

2. single alloyed-single diffused layer,

3. double diffusion technique and

4. lithium ion drift technique.

The last method was first described by Pell (1960) and is the most

popular of the methods. It allows semiconductors of low purity to be

used rather than intrinsic material and provides the capability for

controlling device geometry. By this process lithium is diffused into

p type silicon, until an n-p junction is produced internally. A reverse

bias is then applied to the junction. The electric field in the deple-

tion area will move the positively charged Li+ ions from the lithium

rich side to the lithium deficient side in a reasonable length of time

at sufficiently high temperatures (approximately 1750 C). The amount

of drifted lithium adjusts itself to exactly compensate the negatively

charged acceptors in the bulk material. In this way an intrinsic region

is produced. Blankenship and Borkowski (1962) improved on the lithium

ion drift technique. Using their method a thin dead layer of low

resistance can be formed on the n side of a p-i-n diode. Ammerlaan

and Mulder (1963) reported a very detailed description of the procedure

to be used for preparing p-i-n detectors.

Methods of Detection and Measurement

PN and p-i-n junctions can be used for measurement of radiation in

one of three ways. In order of sensitivity these are pulse counting,

voltage measurement and current measurement.

Pulse counting. In pulse counting a high reverse bias is applied

to enable radiation generated carriers to form a pulse. The amount of

charge collected in each of the pulses is proportional to the energy

deposited in the depletion area. The counting rate depends on the

photon flux and energy, the applied junction potential and the thickness

of the depletion layer. The charge collection time (T) is given by

T = W2/1V


= carrier mobility,

W = depth of depletion area and

V = applied voltage.

The pulse size is not greatly affected by changes in bias or temperature.

The pulses are usually amplified by low noise, charge sensitive


Although pulse counting is the most sensitive of the measurement

methods, it is only useful for dose rates ranging from 10 prad/hour to

100 rad/hour (Baily and Kramer, 1963). Since the dose rates generally

encountered in radiation therapy are higher, this method of operation

would be unsuitable for dosimeters. PN junction detectors which

operate by pulse counting could find useful application in health

physics, when used in conjunction with well designed amplifiers

(Jones, 1962). Baily and Hilbert (1966) reported a study in which

p-i-n junctions were used as pulse counters for y ray dosimetry. Pulse

counting methods have been used for Y rays ranging in energy from 0.1

to 1 MeV (Jones, 1962).

Voltage measurements. Voltage measurements may be used for dose

rates ranging from 1 mrad/hour to 100 rads/min. This method of measure-

ment, which is the most simple, is an analog of the photovoltaic effect

in solar cells. The photovoltaic measurements of p-i-n and pn junctions

used for x and y ray dosimetry are of greater sensitivity than measure-

ments of solar cells (Scharf and Sparrow, 1964). This is due to the

increased resistivity and collecting volume of the dosimeters. In the

photovoltaic mode of operation there is no applied bias. When the

detector is irradiated, the radiation produced carriers in the depletion

region and any minority carriers that reach the junction by diffusion

are acted on by the junction field. A photocurrent is generated in the

reverse direction of the junction as electrons are moved to the n side

and holes to the p side. Thus, a voltage difference is produced between

the opposite sides of the diode. This photovoltage biases the junction

in the forward direction. The induced potential may be read out

directly with a high (10 MQ) impedance voltmeter, i.e., the open circuit

voltage is measured.

The relationship between the voltage and current is given

kT I I
V kT In (g I + 1) (5)
q I


V = open circuit voltage produced by irradiation,

k = Boltzman constant,

T = absolute temperature,

q = electronic charge,

I = generated photocurrent,
I = photovoltaic output current and

I = diode saturation current.
Measurements have been performed to determine the relationship

between open circuit voltage and a number of parameters (Scharf, 1960;

Baily and Kramer, 1963; Scharf and Sparrow, 1964). These parameters

include dose rate, temperature, wavelength, energy, width of the in-

trinsic region and angle of incidence of the radiation. At low dose

rates the voltage response is approximately linear, but at higher dose

rates the response becomes logarithmic. Equation 6 shows the relation-

ship between open current voltage and dose rate.

V 2kT In K D (6)
q 1


K = a constant of the detector and

D = dose rate.

Whelpton and Watson (1963) found a linear response of open circuit
voltage up to a dose rate of 750 R/min of Co y rays in a pn junction,

while Baily and Norman (1963) found linearity up to a dose rate of

400 R/min using a p-i-n diode. The voltage response is highly dependent

on temperature, it decreases nearly exponentially with increasing

temperature. The wavelength dependence of the open circuit voltage is

due to variations in the carrier generation rate. Since the variation

rate is small for energies greater than 200 kev, the voltage

measurements become wavelength independent above this energy (Parker

and Morley, 1966).

Current measurements. If a detector operates in the photovoltaic

mode an alternative to measuring the voltage is to measure the photo-

current generated. As stated previously, the photocurrent generated by

radiation is proportional to the intensity of the radiation absorbed.

It consists of the drift current (I ) of electron-hole pairs produced
g Dr
in the depletion region and the diffusion current (I )D of minority

carriers that reach the depletion area by diffusion. If the load

resistance across the detector terminals is zero, the short circuit

current is equal to these two currents which flow through the external

circuit in the reverse direction. If the load resistance is greater

than zero, there will also be a diffusion current (Ij)Di in the forward

direction. This current is produced through the junction by the photo-

voltage in an attempt to restore the equilibrium state and causes the

total photocurrent to decrease. The net diffusion current will be

(I )Di (I.) .. A sensitive galvanometer is used for measuring
gDi j Di
current. The short circuit current may be determined by extrapolating

photovoltaic current measurements to zero load resistance, but generally

is approximated from measurements at small load resistances.

Several studies have been performed to determine the relationship

between photocurrent and temperature. Scharf and Mohr (1971) found the

temperature dependence of short circuit currents was positive and nearly

linear in pn junctions but nonlinear and negative for p-i-n junctions

and surface barrier type detectors. Klevenhagen (1973, 1977) also

observed positive and negative temperature dependence of short circuit

current and that the magnitude of this dependence varied from detector

to detector. Petushkov and Parker (1973) also noticed anomalies and

suggested that dosimeters should be zeroed at the operating temperature.

Scharf and Mohr (1971) explained that the variation seen by different

investigators of short circuit current with temperature in different

detectors was due to the influence of the strongly temperature dependent

junction current (I.)Di and to the different values of series and

junction resistances. If there is no junction current, the short cir-

cuit current is independent of temperature.

Current measurements may also be performed when an external reverse

bias is applied. This is called the photodiode mode of operation. As

in the photovoltaic mode the current generated by radiation in the

reverse direction of the detector is proportional to dose rate. The

total current measured when the detector is irradiated consists of the

generated photocurrent and the dark current produced by the bias

voltage. If the dark current and total current are measured, the gen-

erated photocurrent can be calculated. The dark current is composed of

thermally induced carriers and surface leakage. As the bias voltage is

decreased the ratio of radiation generated current to leakage current


Studies have shown that photocurrents measured in the photodiode

mode increase with increasing temperature (Scharf and Sparrow, 1966).

The energy dependence is similar to that exhibited in the photovoltaic

mode. Current measurements may be used for dose rates above 1 mR/min.

The choice of a measurement technique will depend on the sensi-

tivity required and the dose rate of the radiation.

Uses of P-I-N Diodes

P-I-N diodes are used mainly as detectors and dosimeters.

Detectors can be made in a variety of shapes and sizes; thus, there is

a high degree of versatility in the measurements that can be obtained.

Baily and Kramer (1963) report that some detectors can be used for

animal and human implantation since they can operate without an applied

voltage or with the exposed portions being held at ground potential.

Baily and Norman (1962) fabricated small cylindrical p-i-n radiation

detectors for internal dose measurements and for use with internally

administered radioisotope tracers. Such detectors can be used to

detect tagged tumor tissues during brain surgery. When these detectors

are used for tracer applications the pulse counting mode of operation is

used and, therefore, a bias voltage has to be applied. A preamplifier,

amplifier and scaler must be used.

The response of p-i-n junctions to f rays emited from 32P and
Th was investigated by Baily and Hilbert (1965). The response which

was determined for various source geometries, detector sizes and energy

distributions of the emitted electrons was compared with the dose rate

measured by extrapolation chamber techniques. West et al. (1962) de-

scribed the reliability they found in using p-i-n detectors as B ray

spectrometers. Greening et al. (1969) used a lithium-drifted silicon

detector for the measurement of energy fluence rates of low-energy

x rays and found that it was as precise as a gas ionization chamber.

However, because the leakage current is greater, the minimum fluence

rates which can be measured is greater. P-I-N detectors are particu-

larly useful for recording pulsed radiation fields.

Besides being used as radiation detectors and dosimeters, p-i-n

junctions have been used as light detectors since silicon diodes are

also sensitive in the optical portion of the spectrum.

Problems Encountered in the Use of P-I-N and PN Junctions

Several precautions must be taken when using p-i-n junctions for

dosimetry. Since diodes are light sensitive, for accurate radiation

dosimetry the junctions should be protected from light. Whelpton and

Watson (1963) used black PVC tape for this purpose, whereas Raju (1966)

coated the diodes with Kodak dull black brushing lacquer.

Bloch and Worrilow (1969) described the use of polyfoam as thermal

shielding for a detector in order to minimize the drift in reverse

current due to thermal changes. Although the radiation induced current

does not vary with temperature, the leakage current varies strongly

with temperature. Drifting of this leakage current will give rise to

errors especially at low dose rates.

In order to prevent electrical leakage from these semiconductor

dosimeters, very thin insulating coatings may be used. Examples of

these insulators are epoxy resins and silicon monoxide.

Wavelength dependence of pn and p-i-n junctions is not a major

problem if high-energy radiation is being measured. Parker and Morley

(1966) found that detectors could be regarded as wavelength independent

for exposure measurements of radiation whose energy spectrum is above

200 kev.

Thomas and Shaw (1978) stressed the need for caution in the use of

diode detectors because of the problems that could result from variation

of response with angular orientation. If the semiconductor material is

not axially mounted in a detector, there may be a directional response.

Semiconductor materials are very sensitive to radiation damage

which produces imperfections in the crystal lattice. Vacancy-

interstitial pairs known as Frenkel defects may be formed as silicon

atoms are displaced from their equilibrium sites. P-I-N diodes have a

thick depletion area and increased sensitivity to radiation and,

therefore, have decreased resistance to radiation damage., In these

diodes radiation damage causes a decrease in charge carrier lifetime

and mobility. Loferski and Rappaport (1958) determined that minority

carrier lifetime (T) was more sensitive to radiation damage than con-

ductivity or mobility. They did direct measurements of T and determined

the minimum energy needed to produce a Frenkel defect. Coleman and

Rodgers (1964) reported that p-i-n junctions are more susceptible to

radiation damage than pn junctions because of the low fields at which

they are operated and the increased collection time and decreased

collection efficiency. The concentration and type of impurity in the

compensated material may also affect the sensitivity of p-i-n junctions

to radiation damage. Rosensweig (1962) noted that when a silicon solar

cell was exposed to electrons and y rays of energy greater than 300 kev

or to heavy particles, radiation damage caused the diffusion length to

decrease. The diffusion length is the average distance travelled by a

minority carrier prior to recombining with a majority carrier. Dearnley

(1964) pointed out that if radiation damage caused a change in detector

response this would lead to a misinterpretation of data.

A final problem to be considered is the energy dependence due to

the difference in Z of silicon and soft tissue. This is important for

low-energy radiation where the photoelectric effect is the principal

mode of interaction.

Despite the problems involved in using semiconductor detectors,

the advantages make them very attractive for in vivo dosimetry. To

reiterate, these include the small size, high sensitivity per unit

volume, the low operating voltage, the mechanical strength, ease of

construction, instantaneous readout and the simplicity of electrical

measuring equipment.

Air Cavities

In vivo dosimetry in radiation therapy is needed in several parts

of the body. If a treatment field is regular, i.e., of normal rectan-

gular shape, the dosimetry procedures are fairly simple and

straightforward. Isodose curves for the given field size are used and

corrections are made for any curvature on the surface. This procedure

may not result in a precise estimation of the dose in a patient because

of the presence of differing types of tissues in the treatment volume.

This problem is compounded even further in irregular fields and fields

treated by rotation therapy.

Many of the fields used when treating head and neck cancer with

external radiation are regular but because of the variation of tissue

across this area the treatment plans may not provide adequate accuracy.

Air cavities, e.g., sinuses and trachea, may play an important role in

decreasing this accuracy. Several studies have been performed to deter-

mine the influence of air cavities on the dose distribution of x and y

rays as well as electrons. One of the older studies was performed by

Epp et al. (1958). They investigated the ionization build-up at the

surface of a lesion located in an air cavity of the upper respiratory
system when irradiated with a Co beam which first passed through the

air cavity. They simulated this condition by measuring the response of

a small flat ionization chamber as the thickness of the front wall was

varied from nearly zero thickness to equilibrium thickness. The con-

clusion was that unless a radiation field was bigger in all directions

than the air cavity by at least the electron range in tissue, the

absorbed dose at the inner surface of the air cavity may be less than

equilibrium value. The dose to a lesion located on the inner surface of

an air cavity may be reduced by as much as 10 percent if electronic

equilibrium is not attained.

Epp et al. (1977) performed similar studies utilizing 10 MeV x rays

from a linear accelerator. They found that the lack of electronic

equilibrium at the surface of an air cavity was more severe for higher

energy rays. Another study of this type was performed by Koskinen and

Spring (1973). They used LiF teflon dosimeters, which were 20-28 pm
thick, to study the dose distribution of .Co radiation in a phantom

simulating the upper respiratory air passages. They devised a formula

for calculating the decrement in dose at the surface of an air cavity.

Nilsson and Schnell (1976) expanded on the previous measurements. They

used 10 pm thick LiF dosimeters to measure the dose in a phantom in

which air cavities of different sizes were placed. The radiation source

used was a 6 MeV linear accelerator and a 42 MeV betatron. Nusslin

(1975) studied the effects of air cavities on the dose distributions of

a high energy electron beam from a betatron, and Samuelsson (1977)

studied the influence of air cavities on the dose distribution of 33 MeV


These types of measurements have particular significance in the

treatment of cancer of the larynx. If a parallel opposed pair treatment

technique is used to treat a laryngeal lesion, one of the beams will

pass through an air cavity before arriving at the lesion, and if adequate

precautions are not taken underdosing may result in the transition zone

between air and tissue. Bagshaw et al. (1965) found that the treatment

plan of choice for most patients with laryngeal cancer was opposed

lateral fields.

Morrison (1971) observed that laryngeal cancer is usually diagnosed

at an early stage and, therefore, has a good chance for cure. However,

Lederman (1971) reported that despite its accessibility and ease of

examination, the radiation treatment of laryngeal cancer has not been as

successful as desired. Most of the tumors are on the vocal cord which

has a poor lymph drainage and, therefore, a low rate of spread to lymph

nodes. Tumors in the supraglottic and subglottic regions have a much

greater tendency to lymphatic spread. With the use of supervoltage
radiation in particular Co, rather than orthovoltage there has been

gradual improvement in cure rate, in addition to the advantage of less

discomfort suffered by the patient and minimizing of reactions.

It is unlikely that there is any advantage to be gained in using

radiation that is more penetrating than 60Co for treatment of laryngeal

cancer. Morrison and Deeley (1962) noted that the only possible ad-

vantage might be a different physical distribution of the radiation. If

8 MeV radiation from a linear accelerator is used, the tissues between

1.3 and 2.1 cm below the skin receive the peak dose, and since this is

the depth of the vocal cord, the dose received would be homogeneous in a

vocal cord lesion and fall off towards the normal tissue.

An analysis of cases of laryngeal cancer which were not cured by

radiation therapy suggested that underdosing may be a major cause of

these failures. This reinforces the need for better dosimetry. A study

performed by Spring et al. (1972) illustrated that there was a correla-

tion between recurrence of laryngeal cancer and the treated volume.

Another study (Salmo et al., 1973) narrowed down the analysis to include

stage 1 patients only. The conclusion was that larger treatment fields

gave a greater probability of cure without recurrence.

A dosimeter which is more accurate than TLDs and smaller than an

ionization chamber would be useful in the continuation of studies of

the effect of air cavities in the treatment with radiation of laryngeal


Objectives of Research

The overall objective of this research was the design and construc-

tion of a miniature dosimeter to be used for measuring doses under

several conditions that are of clinical importance in radiation therapy.

In particular, a study will be made of the effects of air cavities in

the treatment with radiation of head and neck cancers.

At the present time, Nuclear Instruments manufactures a dosimeter

utilizing diodes but such a dosimeter is too large for the application

needed. Basically what is required is a miniaturization of a diode

dosimeter, but with a change in scale has come a change in resources.


In order to determine the feasibility of this project initial

tests were performed on commercial diodes. The diodes used were

RCA 30808, Unitrode UM 9442 and 125-8. The RCA 30808 was an n type

silicon p-i-n photodiode with a 5 mm2 photosensitive area packaged in a

TO-5 can. The Unitrode 9442 is also packaged in a TO-5 can, with the

p-i-n chip bonded to an alumina substrate. The 125-8 was an axially

leaded device. These diodes were exposed to 150 and 250 KV x rays,

filtered with 0.5 mm Cu from a Siemens unit, in order to obtain an

estimate of the current generated .by radiation. The diodes were con-

nected via coaxial cable to a Keithley 616 digital electrometer and

dark current and photocurrent measurements recorded. Of the three

diodes tested the RCA 30808 was the most sensitive and the Unitrode

125-8 least sensitive. A rough check of angular dependence was also

performed. Measurements were taken when the diodes were rotated

through 900 intervals. All three diodes showed some degree of varia-

tion in sensitivity between angles. Table 1 is a summary of the


With the RCA 30808 and the Unitrode 125-8 diodes the maximum dif-

ference in readings at different angles was approximately 22 percent and

with the UM 9442 the maximum difference was 14.5 percent. Such a large

difference would be unacceptable in a dosimeter. Since it would be

Table 1. Summary of Measurements with Commercial Diodes

A e 150 KV, 15 mA 250 KV, 15 mA
Angle Filter 0.5 mm Cu Filter
(degrees) 0.5 mm Cu Filter 0.5 mm Cu Filter

Photocurrent Measurements in nAmp

Unitrode 125-8
Unitrode UM 9442
RCA 30808'



Angular Dependence of Diodes

Unitrode 125-8

Unitrode UM 9442

RCA 30808










NOTE: All measurements were taken using a field
a focus skin distance (FSD) of 60 cm.

size of 6 x 6 cm2 and

extremely difficult to position the dosimeter with the same orientation

each time, a dosimeter is required which shows the same response to

radiation from any direction.

Measurements were made with the p-i-n junction in two different

orientations--parallel to the beam and perpendicular to the beam. It

was noted that the readings were very different in the two configura-

tions, e.g., in the Unitrode 125-8, a measurement of 2.5 nAmps was made

for a given dose when the diode was perpendicular to the beam compared

to 0.462 nAmp when the diode was parallel to the beam.

Another method of dosimetry considered was cadmium telluride

(CdTe). Although CdTe can detect radiation efficiently without the use

of photomultipliers its use has been limited to uptake measurements of

radiopharmaceuticals (Walford and Parker, 1972). Ideally a detector

should have a tissue equivalent response and the main disadvantage of

using CdTe for dosimetry is the large difference between its effective

Z (Z = 50) and that of tissue (Z = 7.8). Cadmium telluride is more

energy dependent than silicon, however, because of its resistance to

physical injury and ultimately increased reliability and its ability to

be fashioned into a miniature detector CdTe may prove to be a good

choice for a dosimeter. The evaluation of a CdTe dosimeter would be a

complete area of investigation and was not pursued in this study.

Description of Ideal Dosimeter

As discussed previously some of the properties of the ideal

dosimeter include

1. small size,

2. lack of angular dependence,

3. lack of temperature dependence,

4. ease of insertion prior to measurement and ease of removal,

5. ease of setting up and taking a measurement and

6. accurate and reproducible measurements.

One of the aims of this research project was to design a dosimeter with

as many of these ideal properties as possible, and which is small

enough for insertion.

The primary limiting factor in constructing the dosimeter as

desired was the lack of sufficient funds. The secondary factors were

the availability of small diodes and coaxial cable. Once the feasi-

bility of using silicon diodes was determined it was essential to find

a manufacturer who could supply extremely small p-i-n chips. Although

several companies manufacture diodes utilizing chips, few were able to

supply individual chips, especially of the small size needed. Unitrode

Corp. could supply chips in two sizes, namely 0.062 in. square and 0.03

in. square.

To overcome the problem of angular and directional dependence the

sensitive area of the silicon detector should have a coaxial geometry as

in Figure 2. This would allow the most compact sensitive volume, and

regardless of the direction of the radiation beam the response would be

the same.

Since p-i-n chips are not available in such a configuration, this

would entail having the chips specially manufactured in this

configuration. However, this would prove extremely expensive and was

beyond the budget of this project; therefore, alternate methods of

achieving this requirement had to be used.





Figure 2. Diagram of Idealized Detector

Since several chips were to be used in the construction of a

dosimeter, the characteristics of each should be checked individually.

Ideally before each chip is used in the dosimeter it should be irradi-

ated and the signal examined. The response to factors such as tempera-

ture should also be determined for each chip. The manufacturer of the

chips suggested that they should be checked for leakage at room tempera-

ture before and after heating them up to 1250 C for 168 hours. Several

chips should be examined and those with similar response used. Several

dosimeters should also be made for testing. Once each dosimeter is

calibrated the same tests should be performed on each to determine if

there is any variation in the response and thereby check the accuracy.

Since the sensitivity of the dosimeter may decrease with cumulative

radiation damage, calibrations should be repeated periodically. The

cost of carrying out all these procedures was prohibitive; therefore,

many of them were omitted.

Since the accuracy needed in radiotherapy is 5 percent, the

dosimeter has to be of even greater accuracy since there are several

other sources of error. The signal to noise ratio should be better

than 100:1. The diode dosimeters that are currently manufactured com-

mercially are intended for use on the surface and utilize a single chip.

In this design in order to have a large signal and equal directional

sensitivity several chips were used.

Prototype 1

In order to simulate a coaxial geometry of p-i-n diode the first

dosimeter constructed was made with three chips positioned on the cir-

cumference of a supporting substrate as shown in Figure 3. The chips





Figure 3. Diagram to Illustrate Arrangement of p-i-n Chips in
Prototype 1

used were 0.062 in. square. The supporting rod was made of aluminum and

served as one of the electrodes. The signal from each chip was con-

nected in parallel. The entire structure was surrounded with potting

compound which supplied mechanical strength and protection and was

placed in polyethylene tubing. The signal was led through a 2 ft long

miniature coaxial cable manufactured by Microtech. The dosimeter was

assembled by Eltec Instruments, Inc. The diameter of this dosimeter

was 3 mm and it was 1.1 cm in length. Kodak dull black lacquer was

used to make the device light-tight. However, initial measurements

indicated that the device was extremely sensitive to light and electri-

cal interference. Since this detector would not be used in patients,

materials were then used which would be undesirable if the dosimeter

were to be used clinically. The dosimeter was painted with Television

Tube Koat and this was connected to ground potential. This served to

make the device insensitive to light and static electrical interference.

This first prototype which is shown in Figure 4 was not enclosed

in a stainless steel tube in order to reduce the number of variables in

the measurement and to better evaluate the performance of the diode.

The tubing would also serve to attenuate the beam and act as build-up.

Several tests were performed to check different physical parameters

and the measuring device used was a Keithley 616 digital electrometer.

The diode was used in the photovoltaic mode with zero bias voltage and

the short circuit current measured. The reproducibility was checked as

was the linearity. The directional sensitivity and temperature depen-

dence were other characteristics of the dosimeter that were measured.

These measurements are of particular importance since the dosimeter is

being designed for in vivo studies. The response to temperature has

Figure 4. Photograph of Prototype 1

" .I .. .".. .'.*"..I *| Ti* s ' I T i iniT I (TITI i T i T i il il
1.. 21 2 -4 -5 6 7 8 9 .''O --1 0 11 1 2 '1 1 4 ""
Sli li 1 i Ii ll it i li i l i i -il i ii il IIij ,il ,l h ,, ,l,, ,IIIn I i n I . I I' J ..i. A .. u... 1 ... ..

been seen to vary for different diodes and should be determined for

each dosimeter. The procedures for and the results of the measurements

are discussed in Chapter 4.

There were several problems that had to be overcome in the con-

struction and use of this dosimeter. Since the dosimeter was intended

for interstitial use, the use of inert non-toxic materials was a re-

quirement that had to be fulfilled. The design called for the dosimeter

to be encased in thin stainless steel hypodermic tubing; however, pro-

visions had to be made in the event that the needle should be ruptured

at its insertion or removal. An inert potting compound was used around

the chips. Another potential problem is the lack of tissue equivalence

of the materials in the dosimeter. It is desirable that the effective

Z of the dosimeter be close to that of tissue; however, since the dosi-

meter will be used at energies where Compton interaction predominates,

this is not an essential factor provided a proper calibration is

obtained. The problem of cumulative radiation damage may be minimized

by regular periodic calibration of the dosimeter. Thus, if there is a

change in sensitivity, a new calibration factor will be established.

The tests of the first prototype indicated there was some amount

of angular dependence. The maximum difference was 12 percent. This

could have been due to the use of a defective p-i-n chip or to the im-

proper placement of the chips. Since it was not possible to check the

response of each chip individually, a different arrangement of the

chips was used in the next dosimeter.

Prototype II

In the second dosimeter the six chips used were each 0.03 in.

square, and again they were connected in parallel electrically; however,

they were placed side by side. Unitrode Corp. (the manufacturers of

the chips) indicated that the chips would not exhibit angular or direc-

tional sensitivity. The dosimeter which is shown in Figure 5 was

encased in 14 gauge hypodermic stainless steel tubing. The criteria

for the choice of the probe material were strength, hygiene and ease of

insertion rather than tissue equivalence. The response of this dosi-

meter is discussed in Chapter 4.

Peripheral Apparatus

The basic requirements for displaying the signal from the dosimeter

are an amplifier with an integrating feedback circuit, a switch for

resetting after each measurement, an analog to digital converter and a

display panel (Figure 6). The current generated in the dosimeter by

radiation is extremely small and should be amplified prior to display

to give statistically significant results.

The actual measurement and amplification of the radiation generated

current was performed with an operational amplifier operating in inte-

gral mode with a feedback capacitor. An operational amplifier (op amp)

consists of several transistor amplifiers connected in series but has

the advantage over transistors of being able to provide a higher gain.

Operational amplifier circuits were originally developed to perform

mathematical operations for analog computers using feedback but now

have numerous applications, among which are instrumentation, use in

Figure 5. Photograph of Prototype 2


i TT!'I' 1 IT"i 11 i!1 ir I1I~ll I iiI 1iir If rIT 1111 111, 11 11,I' I i' ily i i i 1 1VIIIIII I T TI T
mm 1, :-31TK4 -T5 M6 `7 S ,,, 9 ---10 1 1 1,2 1 3 1'4 "I`




Figure 6. Diagram of Requirements for Displaying Signal

control systems and in regulating systems. The main properties of op

amps are

1. infinite voltage gain,

2. zero output impedance,

3. high input impedance,

4. capability of maintaining the above properties over a particu-

lar frequency range,

5. zero voltage across input terminal and

6. absence of changes with environmental conditions, e.g.,


These features are typical of an ideal amplifier. In reality an op amp

does not possess these features but is close to achieving some of them.

A typical op amp will have high gain, high input impedance, low output

impedance and wide bandwidth. Since all parameters cannot be optimized,

the choice of a suitable op amp depends on its application. In this

application the current was integrated over the time of exposure and the

amplifier produced an output voltage which is proportional to the inte-

gral of the input voltage. Figure 7 shows the circuit of a typical

integrator using an op amp. This circuit incorporates a voltage divider

network which will be discussed later.

The gain between input and output is largely independent of the

gain of the amplifier but is determined by other elements in the

circuit. The output of the amplifier is fed back to the negative or

inverting input through the feedback capacitor. Since the rate of

change of output voltage is proportional to the input voltage, the basic

equation that describes the functioning of an integrator is


Figure 7.

Typical Integrating Circuit




V V dt

o RC i


V is the output voltage,

V. is the input voltage and

C is the feedback capacitance.

For this application an op amp designed for electrometer applica-

tions would be especially useful. This circuit should give accurate

results when measuring small currents, i.e., currents in the nanoampere

range. An important op amp parameter for this application is the input

offset voltage. This is the voltage that must be applied across the

inputs of an op amp or to some other point specified by the manufacturer

to ensure that the output is zero when there is no signal input source.

Sufficient voltage can be applied to null out an unwanted signal, e.g.,

the dark current of the dosimeter. Another important parameter is the

input offset current which is the difference between the bias currents

going to the inputs. The circuit used for measuring the signal is

shown in Figure 8.

Amplifier Circuit

Operational Amplifier

In the first circuit an Intersil ICH 8500A amplifier was used.

This was expected to be ideal for this application since it is designed

for measuring currents in the pico-ampere range. However, because of

problems with the offset voltage null the circuit was redesigned and a

Figure 8. Operational Amplifier Integrating Circuitry


;.205 Ka

:.582 KG

:1.52 Ka

7.74 KG

10 KA

Datel AM 490-2 op amp was used. This amplifier is chopper stablizied

and, therefore, has the advantage of having low input offsets and

temperature coefficients. It can also be used successfully for pico-

ampere level signals. The method of zeroing the offset voltage recom-

mended by the manufacturer was used with a few changes. Although the

Datel AM 490-2 showed greater stability than the ICH 8500A, there was

still a problem of reducing the dark current reading by using the input

offset null adjustment. The method for zero adjustment suggested by

the manufacturer is applicable if the inverting input is very close to

ground potential. An offset current is injected into the inverting

input by means of a resistor connected to a potentiometer. However,

the silicon p-i-n chips had a greater resistance than could be success-

fully handled by the circuit.

In the final circuit a Teledyne Philbrick Model 1702 op amp was

used. A 50 KQ potentiometer was used to zero the voltage offset.

Feedback Capacitor

Since the feedback network is the main determinant of the char-

acteristics of the circuit, the type and value of the feedback capacitor

used is important. As shown in the circuit diagram negative feedback

was used, the signal going to the negative input. Previous measurements

with the dosimeter utilizing a Keithley electrometer indicated that the

signal was in the nanoampere range. The time in which integration

would occur would be a typical treatment time of about 200 sec. Since

the current I is 2 x 10-9 amp and the rated amplifier output is T 10 V


dV 10 V 0.05 V/sec
dt 200 sec

But I ,


C = 2 x 109 amp = 0.04 pF.
0.05 V/sec

A low loss capacitor was used. Because the value of the feedback

capacitance is fairly large, the shunt or stray capacitance (Cs) did

not affect the values obtained in further calculations and was omitted.

The capability of adjusting the sensitivity of the amplifier system

was needed. One method of achieving this is by the use of different

capacitors and switching to the one required for a particular

measurement. A simpler method and the one that was used in this device

was a voltage divider. This network served to alter the effective feed-

back capacitance. A simple voltage divider network incorporated in an

integration circuit was shown in Figure 7.

Since I = V (Ohm's law)

and IR I (i.e., the current flowing into R1 and R2

is equal)

V V1 V1
then _(7)
R1 R2

Solving equation (7) for V1, it is seen that

R2 V0
V = R2+ (8)
1 R + R2


If = R2/(R + R)

then V = R V,
1 E 0,

and since

Cd V1
I =

then I = -C R (9)

Solving equation (9) for V leads to


V 1 Idt. (10)


The gain (A) of the voltage divider network is V /V where


V 1 I dt (11)
n C J


RE-I dt

A = (12)

S1 I dt

which ultimately reduces to

C RI + R2
A = -

Thus, by selecting the appropriate values of R1 and R2 the gain and,

therefore, the sensitivity of the amplifier system can be regulated.

The gain becomes unity when R2 = R1 + R2.

The effective feedback capacitance (CE) can be calculated

E = C ( R + R
1 2


In this circuit a fixed feedback capacity of 0.04 pF was used and the

total resistance of the voltage divider network was 20.35 KQ. Table 2

lists the gain characteristics of the system.

Table 2. Gain Characteristics of the Amplifier System

Gain Added Resistance Effective Feedback
Switch R2 (KQ) Each Position A Capacity
Position (KQ) (pF)

7 20.35 9.95 1.0 0.04
6 10.40 7.74 2.0 0.02
5 2.655 1.52 7.7 0.005
4 1.135 0.582 17.9 0.002
3 0.553 0.205 36.8 0.001
2 0.348 0.1 58.4 0.0007
1 0.248 0.248 82.0 0.0005


Reset System

When the input goes to zero the output voltage does not go to zero

but holds the accumulated integral. Before each measurement a means of

discharging the capacitor is essential to establish an initial condition.

A switch is a simple means of achieving this. In the reset mode the

initial condition is established.

d V
d Vo 0

The switch is closed prior to t = 0, just after it is opened the voltage

across the capacitor is V i.e., the value of Vo established in the

reset mode serves as the initial condition for the integration time


An electronic means of switching was used in the first circuit,

namely a CMOS quad bilateral switch. In essence only a single switch

is needed but because of the difference in voltage on either side of the

switch at the input and output of the amplifier and the increased chance

for leakage, a T network was used in which the output of one switch was

connected to the input of the other and this junction connected to

ground potential across a resistor. Thus, two of the switches in the

4066 package were used. When the switch is open during integration the

voltage on the switch is at ground potential. At the time of reset,

15V is applied to the control of the switch which then discharges the

capacitor. This method of switching was not used in the final circuit

because the integrated circuit (IC) was destroyed when the signal was


In the final design an electromechanical method of switching was

used, namely a reed relay. Basically, a reed relay consists of a coil

which supplies a magnetic field to a metallic switch. Since the switch

is made of magnetic material, it is controlled by the coil. The ad-

vantage of using a reed relay switch for the reset is the simplicity

with which it can be placed in the circuit. The type used was IC

compatible with axial leads. If the current being measured was smaller

than about 10-9 amp, this method would be unsuitable since the switching

noise would be too large for the feedback capacitor used.

After an exposure the output of the amplifier should ideally remain

constant until the circuit is reset and the capacitor is discharged,

prior to the next measurement. There are several factors, however,

which will cause this value to change with time. These factors include

amplifier voltage offset, switch leakage, amplifier input leakage,

capacitor leakage and dielectric absorption and most significantly in

this device the dark current of the dosimeter. Therefore, a means of

nulling these effects or keeping them as stable as possible is needed.

If the leakage is stable then it can be subtracted out of the signal.

In the circuit with the AM 490-2 slight changes were made in the values

of the components suggested by the manufacturer for the offset voltage

adjustment to compensate for these other effects.

Circuit Construction

All the components for the circuit were placed on a small printed

circuit board. Additional components not previously mentioned include

capacitors for bypassing the positive and negative voltage supply to

ground. They serve to protect against high frequency noise.

Special precautions were taken with grounding and protection of the

input of the amplifier. In order to prevent ground loops a common

grounding point was used for the amplifier and digital panel meter. A

guard ring was placed around the inverting input of the AM 490-2 and

attached to ground. Grounded shields were used around the signal cable

and the other apparatus. In order to avoid leakage paths at the input

pin, the wires that had to be connected to the input were connected

directly rather than by the circuit board. A socket with Teflon inserts

was used for the operational amplifier.

Since when in use the amplifier is in the radiation therapy treat-

ment room, it is subject to scattered radiation although it is not in

the direct beam. In order to prevent ionization in the air surrounding

the amplifier, thereby increasing the signal, the housing for the cir-

cuit board was filled with Ceresin wax. The housing was simply an

aluminum box connected to ground potential. It served as a shield

against radio frequency (RF) and electromagnetic interference. In order

to make the amplifier system as compact as possible, the power supply

was placed outside the treatment room in the housing with the digital

panel meter (DPM). The input voltages, the output of the amplifier and

the reset signal were led through a shielded cable to the DPM. This

cable could be disconnected from both the amplifier and DPM. A short

miniature coaxial cable carried the signal from the dosimeter.


The analog to digital converter and display panel used was a

Fairchild Model 53 three and one half digit panel meter. This was

housed in a metal cabinet that could be easily positioned for

measurements (Figure 9). The DPM was powered directly by AC voltage.

Since the maximum output voltage of the DPM was 200 mV and the output

of the op amp is approximately 10V, a voltage divider network was used.

As mentioned previously, the power supply for the op amp was also placed

in the cabinet for the DPM. A 1.5 amp fuse was used as a protective

measure in the device. A reset button was placed on the front panel of

the cabinet which on depression would energize the reed relay and dis-

charge the capacitor resulting in a reading of zero. An on-off switch

for regulating the power to the device was also placed on the front


Figure 9. Photograph of Amplifier and Display System

r -.r -~

", -,J "i ----- .........

.:o -. ;or

.- - .S :
*F. --

*'**' I.-

o or
.- -

- Aa

. . -. .

S;4 .. ._


On exposing dosimeter 1 to Y rays from a 6Co treatment unit which

had a dose rate of 107.6 rads per min, a current of 2.95 nAmp was

generated. This was approximately 196 times greater than the dark

current which was 0.015 nAmp. Dosimeter 2 generated a current of 2.19

nAmp when exposed to the y rays and had a dark current of 0.011 nAmp.

Tests on the Dosimeters


The method of checking the reproducibility was to simply take 10

measurements consecutively using the same set-up each time. The maximum

variation between any two readings was 0.56 percent.


Each dosimeter was exposed to increasing amounts of radiation and

the photocurrent measured. The results are shown in Figures 10 and 11.

There was an exact linear relationship between dose and generated photo-

current for all the doses examined.

Temperature Dependence

For the temperature dependence measurements each dosimeter was

exposed to y rays while immersed in a water bath whose temperature

Figure 10. Graph to Demonstrate Linear Response of Dosimeter 1 to Radiation



of Exposure Response

32-4 sec
212 4
242 4

8-12x 108C
15- 7x 108
23-3- IOCF
309x x1-8
383x 10-B
46 Ox 10-8
53-5x 1078
61 -0 10 "

122.4 152.4 1824 212.4 242.4



o 50

o 40
" 30



Figure 11. Graph to Demonstrate Linear Response of Dosimeter 2 to Radiation


50 -


STime of Exposure Response
30 0
o 32-4 sec 63 x 10 C
z 62-4 12-3 x 10C
92-4 18-1 x 10 C

. 20 122-4 24-0x 108C
z 152-4 29-8 x O8 C
3 0 182-4 35-7 x 10- C



324 62.4 92.4 1224 152.4 182.4 212.4

was controlled. The temperature of the water was allowed to increase

gradually from 230 C up to 400 C. For each of the temperatures for

which a measurement was taken, a reasonable length of time was allowed

to lapse between the water attaining the temperature and the time of

measurement. This was to ensure that the dosimeter had attained the

same temperature as its environment. At each temperature the dosimeter

was exposed to the same amount of radiation and several measurements

were taken at each point. The complete experiment was performed several

times to check the reproducibility. It was noted that erroneous results

were obtained if the water was not stirred constantly. Figure 12

demonstrates graphically the results of some of these measurements.

Both dosimeters showed a positive temperature dependence, i.e.,

as the temperature increased the response to radiation also increased.

For a 30 C rise in temperature the reading of dosimeter 2 increased by

approximately 2.2 percent. Because there is a definite reproducible

relationship between temperature and dosimeter response, it was possible

to calculate a series of correction factors, one of which could be

applied depending on the temperature of the dosimeter. Thus, there

would be a factor for room temperature and another for body temperature.

These correction factors for dosimeter 2 are shown in Table 3.

Klevenhagen (1978) found that it was possible to stabilize the

detector response over a wide temperature range by.means of selected

load resistance in the external circuit. This method was not tried.

Figure 12. Temperature Dependence of Dosimeter 2







() 350
w 0


Temp Response
340 25*C 333 mV
280C 341 mV
31C 348 mV
34C 359mV
37C 365 mV
40C 372 mV



25 28 31 34 37 40

Table 3. Correction Factors for Increased Response
of Dosimeter 2 to Temperature

Temperature Correction Factor

250 C 1.0
280 C 0.977
310 C 0.957
340 C 0.928
370 C 0.912
400 C 0.895

Directional Dependence

For these measurements each dosimeter was positioned in a cylin-

drical lucite phantom which could be rotated through 3600. The long

axis of the dosimeter was perpendicular to the axis of the beam. At

450 intervals, the dosimeter was exposed to a set amount of radiation

from a 60Co treatment unit and the response recorded. Measurements were

also taken at 300 intervals. As discussed previously, the variation in

response at different angles for dosimeter 1 was as great as 12 percent,

but this may have been due to a defective p-i-n chip. For dosimeter 2,

the maximum variation was 2 percent. Figure 13 demonstrates graphically

the result of directional sensitivity measurements for both dosimeters

1 and 2, when readings were taken at 450 intervals. Table 4 shows the

results for dosimeter 2 when readings were taken at 300 intervals.

Measurements were also taken to compare the response of the dosi-

meter if the radiation beam were parallel to the long axis as opposed to

perpendicular. In both configurations it was necessary to have the same

amount of build-up, i.e., expose the dosimeter so the p-i-n chips were

Figure 13. Directional Sensitivity of Dosimeters 1 and 2




Table 4. Directional Sensitivity Measurements for
Dosimeter 2


Relative Scale Reading
(Average of three measurements)


at the same depth in a phantom. Since the sensitive area of the dosi-

meter is not at the tip of the needle but is below potting material, it

was necessary to estimate the amount of phantom material which would be

equivalent to.the potting material. This thickness of phantom material

was used above the needle when measurements were taken with the sensi-

tive area perpendicular to the beam. The difference in response was

7.25 percent.

Construction Parameters Affecting Response

The amount of attenuation and the amount of build-up provided by

the encapsulating material were determined indirectly in depth dose

measurements. It was especially important to determine the effect of

the stainless steel tubing of dosimeter 2. The encapsulating material

of dosimeter 2 was seen to be equivalent to about 5 mm of tissue.

The results and their implications are discussed more extensively

in Chapter 5 with the depth dose measurements.

Energy Dependence

In order to determine the energy dependence of dosimeter 2, measure-

ments were performed using the different energies for which the dosimeter

would be used clinically. The energies used were 1.25 MeV Y rays from a

60Co unit and 8 and 17 MeV photons from a Philips SL 75-20 linear

accelerator. The dose rate was 360 rads per min for the 8 MeV photons,

440 rads per min for the 17 MeV photons and 107.6 rads per min for the

6Co source. The dosimeter was also exposed to a radium source whose

average energy is approximately 0.8 MeV. Any differences in the calibra-

tion factors are due mainly to the effect of the encapsulating material

on the dose absorbed rather than energy dependence of the silicon chips.

The actual response of semiconductor materials varies linearly with. the

energy deposited and is independent of the type of radiation which

deposits the energy.

For calibration for the 60Co radiation the miniature dosimeter was

compared with a National Bureau of Standards (NBS) calibrated Vittoreen

R-meter Model 621 using the standard protocol established by SCRAD

(1971). Measurements were taken with the Victoreen at a 5 cm depth in

a water phantom using a 10 x 10 cm field and the dose rate and dose

determined. Then, using the same geometry, measurements were performed

with the miniature dosimeter.

For the 8 and 17 MeV photons, measurements were taken with a

Capintec 0.6 cc air equivalent ionization chamber attached to a Keithley

616 digital electrometer and a Keithley 6169 ion chamber interface.

The Capintec had been previously cross calibrated with the NBS calibrated

Victoreen R-meter. The procedure established by SCRAD was again used

with the appropriate CX factors. The calibration factor for the 2Ra

source was obtained by comparing the reading of the Victoreen Model 131

with that of the miniature dosimeter when exposed under similar

conditions. The following are the calibration factors for the miniature


0.132 for 60Co y rays

0.111 for 8 MeV photons

0.108 for 17 MeV photons

0.135 for 226Ra.

By multiplying the scale reading of the display system by the appro-

priate calibration factor, the dose in rads is obtained. If the factors
are normalized to the Co calibration factor, this would give the

following relationship:

0.8 MeV 1.02

1.25 MeV 1.00

8 MeV 0.856

17 Mev 0.856

Dose Rate Dependence

Measurements were taken with dosimeter 2 to determine if the

response was dose rate dependent. The dosimeter was exposed to the

same given dose of radiation (50 rads) at different dose rate settings.

Both 8 MeV and 17 MeV photons from the Philips SL 75-20 linear accelera-

tor were used. The dose rate of the 8 MeV photons was varied from 140

rads per min up to 420 rads per min, and the dose rate of the 17 MeV

photons was changed from 100 to 300 rads per minute. The results are

shown in Table 5. Although there was an increase in scale reading with

increased dose rate, the change was minimal, for 17 MeV it was 1.1


Table 5. Dose Rate Dependence of Dosimeter 2

Dose Rate
Ds R e Relative Scale Reading
radss per minute)

8 MeV Photons

140 440
210 443
290 444
360 445
420 446

17 MeV Photons

100 450
140 452
190 454
240 454
300 455

Tests on Peripheral Equipment

Cable Effect

Tests were performed to determine if there was a change in the

dosimeter response when the cable was irradiated. This change would be

analogous to the "stem effect" seen in ionization chambers. The dosi-

meter was irradiated using a series of field sizes that ranged from
2 2
4 x 4 cm to 4 x 20 cm At each field size two sets of measurements

were taken, one with the cable in the beam and the other with the cable

outside of the beam. The response was the same whether or not the

cable was irradiated.

Position of Amplifier

An amplifier system should ideally be as close as possible to the

detector to eliminate noise and degradation of the signal. However,

it should not be in the path of the radiation beam. Measurements were

taken to determine how much effect the position of the amplifier had on

the dosimeter reading. Measurements were taken first with the amplifier

12 in. from the dosimeter and then with the amplifier outside of the

treatment room, the signal being carried to the amplifier through a

100 ft cable. There was no change in the response.

Gain Settings

Measurements were taken at the different voltage divider gain

settings on the amplifier to check that the gain that was calculated was

actually obtained. The dosimeters were exposed to a given dose at dif-

ferent settings and the readings compared. The calculated and measured

values agreed within 1 percent.


The problems in dosimetry and treatment planning caused by the

presence of air cavities in the treatment volume have been discussed by

several workers. These problems were discussed as they related to
treatment with Co (Epp et al., 1958; Koskinen and Spring, 1973), with

10 MeV photons (Epp et al., 1977) and with 33 MeV photons (Samuelsson,

1977). The dosimeters used in these studies were thin walled ionization

chambers and TLDs. In this study in addition to 6Co y rays, 8 and 17

MeV photons from a linear accelerator were used as the radiation source.

Both dosimeters used were made from silicon p-i-n chips, but dosimeter 2

was enclosed in a stainless steel needle whereas dosimeter 1 was not.

These dosimeters had the advantage of providing immediate readout. There

was no waiting period between measurements. As soon as a measurement

was taken, the diode dosimeter could be repositioned for additional

measurements. It was fairly simple to position the dosimeters in the

phantoms for the measurements. When TLDs are used they have to be

handled with forceps, and if they are paired it is necessary to know and

maintain the identity of each pair. Sievert chambers should only be

exposed to radiation a maximum of once daily. These problems are elimi-

nated with the use of the silicon diode dosimeters.

Depth Dose Measurements

The first set of measurements taken in this series were "build-up"

measurements, i.e., the miniature dosimeters were used to measure depth

doses. Using a 10 x 10 cm2 field size and a SSD of 80 cm, central axis

depth doses were measured for 6Co, for 8 and 17 MeV photon measurements

a FSD of 100 cm was used. Depth dose measurements were also taken for a

6 x 6 cm2 field. The phantom used for these measurements consisted of

6 x 6 in.2 slabs of polystyrene, each slab was 1/16th in. thick. Full

backscatter was provided for all the measurements and there was phantom

material on either side of the dosimeter to eliminate air spaces.

Although it is desirable when measuring depth doses on an x ray machine

to determine the dose at a depth and at the reference point simultane-

ously, this was not done in these measurements.

The results of the depth dose measurements are shown in Figures 14-

19. The point of maximum electronic build-up (Dmax) is 0.5 cm for 60Co,

2 cm for 8 MeV photons from the SL 75-20 linear accelerator and 3 cm

for 17 MeV photons from the same machine. These values are obtained if

measurements are taken with an extrapolation chamber or build-up chamber.

In these miniature silicon dosimeters the materials surrounding the

p-i-n chips, namely the potting epoxy, polyethylene tubing and stainless

steel tubing, provide build-up and reduce the thickness of phantom

material that must be used to give full build-up. From the measurements

with dosimeter 2 on the Co unit it was apparent that full build-up was

provided by the encapsulating material. The readings decreased steadily

as the depth of the dosimeter in the phantom increased, indicating that

the encapsulating material was equivalent to at least 5 mm of tissue.

With dosimeter 1, however, the readings increased rapidly at first and

Figure 14. Depth Dose Measurements for 6Co






L 450

, 430


w 420 "



405 Field Size 6x6cm2
SSD 80cm
395 S

0 .79 1.58 2.37 3.16 3.95


Figure 15. Depth Dose Measurements for 60C
Figure 15. Depth Dose Measurements for Co


















.79 1.58 2.37 3.16


Field Size IOx 10cm
SSD 80cm


Figure 16. Depth Dose Measurements for 60
Figure 16. Depth Dose Measurements for Co







430 *

z 425


0 410

w 405


W 395


385 Field Size IOx 10 cm
SSD 80cm


370 i
0 .79 1.58 237 3.16 3.95


Figure 17. Depth Dose Measurements for 8 MeV Photons

260 -

250 -

o 240-
S230 -
^ 220

J 210
C 200
> 190

_j 180 -
170 FIELD SIZE I0 x 10 cm
FSD 100 cm

0 .79 1.58 2.37 3.16 3.95

Figure 18. Depth Dose Measurements for 8 MeV Photons


225 -
220 -
215 -
210 -

I 205
^ 200

q 190
> 185
5 180
175 FIELD SIZE 10 x 10 cm
0 FSD 100 cm


0 .79 1.58 2.37 3.16 3.95

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