X-ray Production
Mohammad Reza AY, PhD
Department of Medical Physics, Tehran University of Medical Sciences, Tehran, Iran
Division of Nuclear Medicine, Geneva University Hospital, Geneva, Switzerland
X-rays – the Basic Radiological Tool
Roentgen’s experimental apparatus (Crookes
tube) that led to the discovery of the new
radiation on 8 Nov. 1895 – he demonstrated
that the radiation was not due to charged
particles, but due to an as yet unknown
source, hence “x” radiation or “x-rays”
Known as “the radiograph of Bera
Roentgen’s hand” taken 22 Dec.
1895
2
1
Chapter 5 Lecture Objectives
v How
x-rays are produced, what spectrum results and
what radiographic technique factors affect the spectrum?
v What elements comprise an x-ray tube and how they
work together to generate x-rays?
v How are x-rays collimated and the exposure timed?
v What is an x-ray generator, how does it assist in the
production of x-rays and how does its design affect the
resulting output spectrum?
v How does the x-ray tube heat loading and cooling affect
the duration and number of radiographic exposures?
3
Photon Interaction with matter
Compton Scattering
Think of the Compton Effect as the way
gamma radiation creates ionization.
Gamma radiation, in the form of a
medium energy photon, hits an atom.
Part of the photon's energy kicks one of
the atom's electrons free. The gamma
photon continues on but at a deflected
angle and lower energy. This continuing
photon is referred to as "secondary" or
"incident" gamma radiation. The freed
electron is called a Compton Electron.
2
Photon Interaction with matter
Photoelectric
In the Compton Effect, part of a colliding
photon's energy is absorbed by an atom,
resulting in a freed electron and a secondary
gamma ray of lower energy.
In comparison, with the Photoelectric Effect, a
photon, usually of a lower energy, collides
with an atom and all of the energy is
absorbed by the atom. An electron is kicked
out and there is no secondary gamma
radiation. This electron is called a
Photoelectron.
Photon Interaction with matter
Rayleigh Scattering
In rayleigh scattering or Coherent scattering
the incident photon interacts with and excites
the total atom. This interaction occurs mainly
with very low energy photons around 15 to 30
KeV. During the coherent scattering event,
the electric field of the incident photon’s
electromagnetic wave expends energy,
causing all the electrons in the atom to
oscillate in phase. The atom’s electrons
immediately radiate this energy, emitting a
photon of the same energy but in slightly
different direction.
3
Photon Interaction with matter
Pair Production
The third way gamma radiation can interact with matter is
Pair Production. This occurs when a high-energy gamma
photon collides with an atom near its nucleus. The
photon's energy is totally absorbed by the atom and two
beta particles are kicked out; one positive (positron) and
one negative (negatron).
The negatron travels through the surrounding matter
creating ionizing atoms along its path until it's incorporated
into another atom or becomes a free electron.
The positron, however, almost instantly collides with a
nearby electron. This results in the annihilation of both
positron and electron and emission of 2 gamma rays of
511 KeV travelling at 180°from each other.
These resulting 511 KeV gamma rays are used in Positron
Emission Tomography (PET) imaging systems.
The Bremsstrahlung Process
Creates a
polychromatic
spectrum
-10
Atom diameter ≈ 10 m
-14
Nucleus diameter ≈ 10 m
12
Volume Ratio ≈ 1:10
8
c.f.: Bushberg,
Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 99.
4
The Bremsstrahlung Process (1)
v X-rays
are produced by the conversion of e- KE into EM
radiation - Bremsstrahlung (G: “braking radiation”)
v A large potential difference is applied across the two
electrodes in an evacuated envelope
v
v
Neg. charged electrode (cathode): source of ePos. charged electrode (anode): target of e-
9
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 98.
The Bremsstrahlung Process (2)
ve
released from the cathode are accelerated towards the
anode with a gain in KE as the e drops through the applied
potential difference (kilovoltage potential - kVp)
v
v
About 99% of the KE converted to heat via collision-like
interactions
About 1% of the KE converted into x-rays via strong Coulomb
(electrostatic) interactions → Bremsstrahlung
10
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 98.
5
The Bremsstrahlung Process
The peak voltage (kVp) applied across the electrodes of the x-ray
tube determines the highest x-ray E (Emax)
The lowest E of the unfiltered x-ray spectrum is not easily
determined, due to severe attenuation of these photons by the
material and thickness of the x-ray tube envelope
v X-ray
production efficiency is influenced by the target Z and
acceleration potential (kVp
(kVp))
11
The Bremsstrahlung Process
X-ray efficiency ≈ Emax · Z · 10-6
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 99.
6
Characteristic X-ray Spectrum (1)
v e-
of the target atom have a
binding energy (BE) that
depends on atomic Z (rem:
BEK ∝ Z2) and the shell (BEK
> BEL > BEM > ... )
v When e-(KE) incident on the
target exceeds the target
atom e-(BE), it’s
energetically possible for a
collisional interaction to eject
the bound electron and ionize
the atom
c.f.: Bushberg, et al., The Essential
Physics of Medical Imaging, 2nd ed., p.
101.
13
Characteristic X-ray Spectrum (2)
v Unfilled
shell energetically
unstable - an outer shell e- with
lesser BE fills vacancy
v As e- transitions to a lower E
state, the excess E can be
released as a characteristic x-ray
photon with E equal to the
difference between the BE of the
e- shells
v As BE are unique to a given
element (Z), the emitted x-rays
have discrete energies
characteristic of that element
14
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 101.
7
Characteristic X-ray Spectrum (3)
v Within
each shell (other than K)
there are discrete E orbitals (ℓ = 0,
1, ... , n-1) → characteristic x-ray
fine E splitting
v Characteristic x-rays other than
those generated through K-shell
transitions are unimportant in Dx
imaging → almost entirely
attenuated by the x-ray tube
window or added filtration
c.f.: Bushberg, et al., The Essential
Physics of Medical Imaging, 2nd ed., p.
102.
c.f.: Bushberg, et al., The Essential
Physics of Medical Imaging, 2nd ed., p.
101.
15
THE X-RAY TUBE
Motor
Stator
Rotor
Anode
Vacuum
Glass envelop
High Voltage
connections
Casing
Bearings
Oil
Output window
Cathode
The tube is inserted inside a casing and immersed in oil for electrical
insulation and cooling
The casing also shields x-rays emitted in all directions
8
X-ray Tube Cathode
e- source - a helical
tungsten wire filament surrounded
by a focusing cup
v Filament circuit: 10V, 7A
v Electrical resistance heats the
filament and releases e- via
thermionic emission (also lights
up – incandescence – light bulb)
v Filament current adj. controls
tube current (rate of e- flow from
cathode to anode - mA)
v Cathode:
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., pp. 104-105.
X-ray Tube Cathode: Focusing Cup
e- distribution when at
same V as filament (unbiased)
v Isolation from filament and
application of a negative bias V
constrains e- distribution further
(biased)
v Focusing cup slot width
determines the focal spot width
v Filament length determines focal
spot length
v Small and large focal spot
filaments (usu. 0.6 and 1.2 mm)
v Shapes
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., pp. 104.
9
X-ray Tube Cathode: Space Charge Cloud
v Filament
current (A) → filament temperature (T) → thermionic emission rate
v When kVp = 0 an e- cloud (space charge cloud) forms around filament
v Space charge cloud shields the electric field for tube voltages of ≤ 40 kVp → only
some e- are accelerated towards the anode: space charge limited
v ≥ 40 kVp the space charge cloud effect overcome by kVp applied and tube current
(mA) limited only by the emission of e- from the filament: emission-limited operation
v Tube current about 5-10 times less than the filament current in the emission-limited
range
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., pp. 105.
X-ray Tube Anode Configuration
v Tungsten
anode disk
Mo and Rh for mammography
v Stator
and rotor make up the induction
motor
v Rotation speeds
Low: 3,000 – 3,600 rpm
High: 9,000 – 10,000 rpm
v Molybdenum
stem (poor heat conductor)
connects rotor with anode to reduce heat
transfer to rotor bearings
v Anode cooled through radiative
transmission (Stefan-Boltzmann law:
radiance ∝ T4)
v Focal track area (spreads heat out over
larger area than stationary anode
configuration)
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 107.
10
CASING & COOLING
• Non circulating oil, natural air convection
• Non circulating oil, forced air cooling (fan)
• Non circulating oil, circulation of water, air cooled
• Circulating oil and oil-air heat exchanger (CT x-ray tubes)
Anode Angle and Focal Spot Size
Anode
kV
mA
Actual Focal
Spot
Filament
Heating
current
Cathode
Optical focus
Apparent
Focal Spot
Useful
X-ray beam cone
The slope of the anode target allows a larger
area to be heated while keeping the apparent
area from which x-rays are produced as small as
possible
11
Anode Angle and Focal Spot Size (1)
v Anode angle (θ
θ): angle of the
target surface with central axis of
the x-ray output field
v θ range: 7°- 20°
v Why are anodes beveled?
v 1. Line focus principle
(foreshortening of focal spot
length)
v
v
“Effective” focal spot size = length
and width of the focal spot
projected along the central axis of
the x-ray field
Effective focal length = actual focal
length · sin(θ
θ)
v
v
x
θ
x·cos(θ)
sin(0°) = 0, sin(30°) = 0.5
For small angles (< 30°):
v
sin(θ
θ) ≈ (θ
θ/57°)
x·sin(θ)
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 108-109.
Anode Angle and Focal Spot Size (2)
c.f.: Bushberg,
Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 108108-109.
12
Anode Heel Effect
v Reduction
of x-ray beam
intensity towards the anode
side of the x-ray field
v Although x-rays generated
isotropically (4π steradians)
Self-filtration by the anode
and the anode bevel causes
Greater intensity on the
cathode side of the x-ray
field
v Can
use to advantage, e.g.,
PA chest exposure
Orient chest to anode side
Abdomen to cathode side
v Less
pronounced as SID ↑
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 112.
Anode Heel Effect
%
slope
120
Anode
100
80
Filament
60
40
20
+20°
+16°
+8°
+12°
0°
+4°
-4°
-8°
-12°
-16°
-20°
0
Off axis relative dose distribution
for a 20°anode slope
On the anode side:
On the cathode side:
Smaller apparent focus
and lower dose rate
Larger apparent focus
and higher dose rate
Apparent focus size and dose distribution vary with the position in the field
This effect limits the useful maximum field size as a function of the anode slope …
… but is used in some examinations to compensate for body thickness variations
( breast, thorax,… )
13
X-ray Filtration
Filtration: x-ray attenuation as
beam passes through a layer of
material
v Inherent (glass or metal insert
at x-ray tube port) and added
filtration (sheets of metal
intentionally placed in the beam)
v Added filtration absorb lowenergy x-rays and reduce patient
dose (↑ beam quality)
v HVL – half value layer (mm Al)
v
c.f.: Curry, et al., Christensen’s Physics of Diagnostic Radiology, 4th ed., pp. 89, 91.
Voltage Transformation
v A time varying V (→ timevarying I) through the primary
winding creates a time-varying B
v If sinusoidal, then
Vp(t) = Vp·sin(2πft) and
B(t) = B·sin(2πft)
v If
the time-varying B lines are
channeled through a
ferromagnetic core, then a timevarying V is induced in the
secondary winding:
Vs(t) = Vs·sin(2πft)
v Magnitudes
of Vp and Vs
depend on the ratio of the
number of primary (Np) and
secondary (Ns) transformer
windings
Rem: f = 1/T
sin(2πft) =
sin(2πt/T) =
sin(360°· t/T)
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 117.
14
Transformer Relationships
v Law
of transformers:
Vp / Vs = Np / Ns or
Ns = Np · (Vs / Vp)
v Step-up
transformer:
Ns > Np
v Isolation
transformer:
Ns = Np
v Step-down
transformer:
Ns < Np
v Equality
v
of power output:
Vp · Ip = Vs · Is
c.f. Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 118.
Autotransformer
v An
iron core wrapped with a
single wire
v Self-induction rather than
mutual induction
v Conducting taps allow the
input to output turns to vary,
resulting in small incremental
change between input and
output voltages
v A switching autotransformer
allows a greater range of
input to output values
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 118.
15
X-ray Generator Components
v High-Voltage
v
v
v
power circuit
Low input voltage
High output voltage
Autotransformer allows kVp
selection
v Filament
v
v
v
circuit
mA sets the tube current
sec sets the exposure
duration
manual exposure or
phototimed
c.f. Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 123.
Operator Console (Technologist)
v The operator selects the peak kilovoltage (kVp), the tube
current (mA), the exposure time (sec) and focal spot size
v The kVp determines the x-ray beam quality (penetrability)
which plays a role in subject contrast
v The x-ray tube current (mA) determines the x-ray fluence
rate (photons/cm2-sec) emitted by the x-ray tube at a given
kVp
v mAs = mA · sec (exposure time) ∝ photons/cm2 (fluence)
v Low mA selections allow the small focal spot size to be
used and higher mA settings require the use of large focal
spot size due to anode heating considerations
16
SINGLE PHASE GENERATOR
HV Transformer
AC
Rectifier
Historical
Not sold anymore
HV
DC
HV AC
X-ray
Tube
kVp = 1.4 average kV
50Hz
2 pulses per period
X-ray energy
Low cost but:
• 100 % ripple
• production of undesirable soft X-rays
• kV control by steps (stepping transformer)
• require high power mains outlet
• bulky
THREE PHASES GENERATORS
HV Transformer
AC
Rectifier
HV AC
HV
DC
X-ray
Tube
6 pulse
Ripple = 13%
6 pulses per period
X-ray energy
50Hz
12 pulse
Ripple = 3.5%
12 pulses per period
17
HIGH FREQUENCY GENERATOR
Rectifier
Chopper
HV Transformer Rectifier
Tube
kV control closed loop
AC
DC
HF
HV
DC
X-ray
Tube
25-100 kHz
X-ray energy
• Very few soft X-rays (fast ramp up)
• Accurate and flexible kV adjustment
• Small size
• Still require high power mains outlet
Power and energy (W and J) should
always be calculated using kVmean,
whatever the ripple is.
kVmean and kVp
… a confusing story!
kVp = 1.4 kV
Single
phase
kV
Ripple = 100%
kV
3 phase
6 pulses
kVp = 1.01 kV
Single phase:
the large difference lead to define the HU
for energy (using kVp): 1J = 1.4 HU
6 pulses:
the 5% over calculation leads to say
1J (calculated with kVp) = 1.35 HU
Ripple = 3.5%
kVp = kV
HF
Practically, they are often expressed
using kVp
Ripple = 13%
kV
3 phase
12 pulses
kVp = 1.05 kV
P (W) = kV x mA
E (J) = P x s = kV x mA x s
12 pulses or HF:
the difference is not significant.
Ripple = 0%
18
BATTERY POWERED GENERATOR
Charger Battery Chopper HV Transformer Rectifier Tube
kV control closed loop
Charge / Usage
AC
HF
HV
DC
Battery
X-ray
Tube
A rechargeable battery stores the electrical energy
• Used for Mobile systems
• Same performances as fixed HF generators (but lower power)
• Large autonomy (20 000 mAs)
• Off-operation charging from a low power mains outlet
CAPACITOR DISCHARGE GENERATOR
Charger Capacitor Chopper HV Transformer Rectifier Tube
kV control closed loop
AC
HF
HV
DC
Small capacity
energy storage
(capacitor)
X-ray
Tube
A capacitor provides the peak power and smooth
the power consumption
• Used for Mobile systems
• Very short autonomy (ex: 50 mAs @ 100kV)
• practically needs mains connection, but low power outlet
• Auto re-charge, wait time indicator
• Capable of very short pulses (1 ms)
• no kV close loop: non constant output (falling load)
• Small and light weight
• Low cost
19
EXPOSURE SWITCHING
During ramp up or ramp down, soft X-rays are
produced, not used for imaging and harmful for
the patient
X-ray energy
Ramp up and down times prevent also fast image
repetition rate
Time
3
1
Primary switching: AC level
the simplest but the slowest
2
Secondary switching: HV DC level
faster, necessary for Angio
3
1
HV
transformer
and
rectifier
X-ray
Tube
2
Grid switching
the fastest, used in Cine mode
Generator Circuit Designs Single-phase
(Half-wave & Full-wave) Rectifier Circuit
Diodes – either vacuum tube or solid-state
device: e- flow in only a single direction (cathode to
anode only)
v
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 125.
20
Complete Single-Phase Two-Pulse Rectifier Circuit
high
voltage, low
current
low voltage, high
current
c.f.: Bushberg,
Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 126.
Single-Phase and Three-Phase Generators
Single-phase generator
Tube current for specific
filament current non-linear
below 40 keV due to space
charge effect: inefficient and
contributes to patient dose
Cable capacitance smoothes
Minimum exposure time =
1/120th sec
Three-phase generator
v
v
v
v
Three single phase
waveforms
Out of phase by 120 degrees
Higher effective voltage
Greater control over
exposure timing
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., pp. 127-128.
21
Phototimers
v Although
a technologist can
manually time the x-ray exposure
(set filament mA and exposure
time or the mAs), phototimers
help provide a consistent
exposure to the image receptor
v Ionization chambers produce a
current that induces a voltage
difference in an electronic circuit
v Tech chooses kVp; the x-ray
tube current terminated when this
voltage equals a reference
voltage
v Phototimers are set for only a
limited number of exposure levels,
thus +/- settings
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 134.
Factors Affecting X-ray Emission
v Quantity
= number of x-rays
in beam
2
∝ Ztarget · (kVp) · mAs
v Quality
= penetrability of xray beam and depends on:
kVp
generator waveform (%VR)
tube filtration (mm Al)
v Exposure
depends on both
quantity and quality
v
Equal transmitted exposure:
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., pp. 136 and 137.
22
Generator Power Ratings and
X-ray Tube Focal Spots
v Power (kW) = 100 kVp · Amax (for a
0.1 second exposure)
100 kW = 100 kVp · 1000 mA @
100 ms exposure
v Amax limited by the focal spot: ↑ focal
spot → ↑ power rating
v Generally range between 10 kW to
150 kW
v Typical focal spots
v
Radiography: 0.6 and 1.2 mm
v
Mammography: 0.1 and 0.3 mm
c.f., Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 139.
X-ray Tube Heat Loading
v Heat
Unit (HU)
HU = kVp · mA · sec · factor
HU = kVp · mAs · factor
factor = 1.00 for single-phase generator
factor = 1.35 for three-phase and high-frequency
generators
factor = 1.40 for constant potential generator
v Heat
input (HU) ≈ 1.4 Heat input (J)
23
Anode Heat Input and Cooling Chart
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 142.
Housing Cooling Chart
c.f.: Bushberg, et al., The Essential Physics of Medical Imaging, 2nd ed., p. 144.
24
Thanks For Your Attention
25