1
12.2 Thermography
Many pathological processes in the human body are closely connected with local or
general temperature change. In section 11.2 we considered thermal measurements at only one
location at a time. This can be made rather precisely. In some cases, the temperature
distribution over the body surface is also of medical interest. Such continuous mapping of
thermal fields is known as thermography. The measuring detector may or may not be in
contact with the body.
12.2.1 Contact thermography
This method of temperature mapping is based on application of cholesteric liquid
crystals. These crystals have thermooptical properties consisting of changes of spatial
arrangement of liquid crystal molecules depending on the temperature. The axes of liquid
crystal molecules form a spiral-like structure. By increasing temperature, the distance between
molecules decreases and vice versa. In this way, the steepness (slope) of the spiral as well as
the wavelength of reflected light decreases. At the lowest temperature the colour image is
characterised by the longest wavelength (red colour), by increasing temperature the colour
image is shifted to shorter wavelengths (yellow - green - blue). Every hue corresponds to a
temperature and represents the isotherm – a region of identical temperature. In medical
practice, the contact thermography is now replaced by the contactless method.
12.2.2 The contactless thermography
The principle of the contactless thermography is detection of infrared radiation (IR)
emitted from the body surface. Spectrum of the human body radiation is relatively broad with
a maximum at about 9 m and respective energy of 0,1 eV. The recording is done from
distance by means of optoelectronic systems which transform the energetic IR map of the
selected part of body surface in a grey or colour scale image. These systems are also denoted
as thermovision.
The thermographic system consists of four main parts: an optical system (objective),
radiation detector, signal processing unit, and a display plus electronic memory (Fig. 12. 2.
2a). The IR image recording is very similar to the recording of visible image by means of a
digital camera. First, the IR radiation propagates through the air, then is directed by objective
lenses, filtered, and produces an image on the detector. Normal glass can be used only for IR
light of very short wavelength which is emitted by very hot bodies. Therefore, the objective
lenses are made of germanium or ZnSe.
The detector used in the IR camera is a transducer which transforms the IR radiation
energy into electric signals. We can distinguish quantum detectors (so-called photonic, made
of InSb or PbSe, for example). They are based on the photoelectric effect. They are accurate,
very sensitive, and produce thermograms with high resolution (1344×784 pixels and more).
However, the photonic detectors have these excellent properties only at low operational
temperatures, and must be cooled, mostly by the Peltier effect. A little less accurate are the
microbolometers which principle is close to thermistor principle. There is no need of cooling,
are much cheaper but their sensitivity is not so high. These detectors have also problems with
thermal stability. Their resolution is smaller (often only 320×240 pixels). Both quantum and
microbolometric elements are organised in a matrix of detectors. The electric signals
produced in detectors are consequently processed and displayed.
2
The display shows the temperature maps (thermograms) with different speed depending on
the kind of the detector. In the quantum detectors it can be even thousands of images per
second, in the microbolometric detectors it is only 30-60 images per second.
In the grey scale thermogram, the cold areas are dark, and the warm areas are white (Fig.
12.2.2b). In a coloured thermogram, each temperature is coded by certain colour which also
represents an isotherm. The modern devices produce mostly the colour thermograms. They
are able to show temperature profiles of the examined areas and perform simple image
analysis. The recorded thermograms can be processed in the same way as any other image
information by means of a computer software on-line or off-line (e.g., determination of the
temperature of any depicted body surface point, calculation of areas of isotherms, different
colour palette settings, adjustment to external conditions and statistical evaluation of the
results).
A factor which complicates accurate measurement of the surface temperature by means of
IR cameras is the surface emissivity or the so-called reflected temperature. The external
medium and close heat sources can be mirrored by the surface of the examined object and
thus influence the measured temperature of the object (the same laws of reflection and
refraction are valid for infrared light as for the visible light). Emissivity is defined as the ratio
of the radiant power of the absolutely black body radiating at given temperature, and the
radiant power of the measured body. The infrared cameras are calibrated by means of the socalled
black body since the black bodies have emissivity close to 1,000. It means that their
“emitted” temperature is not influenced by the reflected radiation. Typical values of human
skin emissivity (independent on skin colour) is 0,96 - 0,98. This value is valid for the skin
without hairs, dry and without any bandages, ointments, lotions etc., which can diminish the
emissivity values. In examination of a patient we also must consider the parameters of
surrounding medium (optimum room temperature, humidity) and also long enough
acclimatization of the patient or respective body surface part before examination.
Obr. 12.2.2a Scheme of IR radiation detection by means of a thermocamera
3
Obr. 12.2.2b Grey scale thermogram of a face. Temperature scale on the right side.
Diagnostic value of thermography
About fifty years ago, thermography was connected with great expectations in noninvasive
screening of some oncologic diseases, namely breast cancer. However, this hope was
not lived up to. Today we use thermography rather as a complementary imaging method in
diseases which are connected with temperature changes of body surface also called dynamic
thermography. We can admit that a substantial decrease of prices of thermographic
equipment, higher resolution and detector sensitivity, as well as a possibility to record video
sequences made this method attractive again for clinical medicine which is demonstrated
mainly in the clinical research up to now.
The sensitivity of modern thermographic systems to temperature changes achieves about
0.025 K. However, in medical imaging, the sensitivity is influenced by localisation of heat
sources. It is relatively high in superficial lesions or lesions just beneath the skin. The deeper
is the lesion the lower the sensitivity. It is caused by different thermal conductivity of tissues
between the skin area heated and the heat producing lesion. These tissues can carry the heat
away or influence its transport in some other way so that the lesion projection on the body
surface need not necessarily correspond to its localisation in depth.
Specificity of thermographic examination is low, in general, as we detect only the place
with the increased or lowered temperature, which can be caused by various pathologic
conditions – an inflammation, tumour, metabolic disorder, or abnormal blood supply. In most
cases thermography itself is not able to differentiate between these causes. In spite of that the
list of medical indications of thermographic examinations is quite extensive: ischemia of
extremities (Fig. 12.2.2c), impairment of nerves, diabetic neuropathy and angiopathy, tissue
healing after reconstructive surgery, inflammations of musculoskeletal system, or
psychosomatic disorders. In clinical experiments it can be used for control of hyperthermia
and cauterisation, some problems in sports and occupational medicine, in physiotherapy and
also in fast fever screening of people in hospitals or airports or during epidemies. There is a
lot of technical applications of thermography, which can have indirect importance for medical
facilities for example checking heat lines and sources, electric power technologies,
transformers, heaters, thermal leaks from buildings etc.
4
Fig. 12.2.2c Left: acute lower extremity ischaemia. Right: left upper extremity with n. ulnaris paresis,
see lowered temperature of the little finger.
5
12.3 Ultrasound imaging and Doppler methods
12.3.1 Theoretical basis
Ultrasound is a mechanical wave motion, the frequencies of which are above the level of
human hearing (20 000 Hz). Most diagnostic applications of ultrasound use frequencies in the
range of 2 - 20 MHz. The energy of ultrasound travels through the medium as longitudinal,
transverse and surface waves. Longitudinal waves are preferred for almost all biomedical
applications. In such waves, the particles of the medium vibrate backwards and forwards
about their mean position. This means that the energy is transferred parallel to the direction of
wave propagation. An ultrasound wave in a given medium has a frequency f, wavelength 
and a propagation speed c, which are related by the equation:
f
c
= (1)
The velocity of ultrasound passing through a liquid medium (in principle, most biological
tissues) depends on the medium's elasticity and density. The velocity in most soft biological
tissues varies between 1500 and 1600 m/s (velocity in water is approximately 1490 m/s).
Some characteristics of the transmission of ultrasound are similar to those of light,
especially when the wavelength is short enough. The ultrasound wave can be directed as a
beam with little scattering and can be focused by an acoustic lens. The ultrasound beam can
be reflected and refracted on the interface of the media and is absorbed and scattered in the
media. All these transmission properties depend on the acoustic impedance of the medium,
which is defined as the product of the ultrasound velocity and the density of the medium. The
characteristic acoustic impedances Z of some materials are given in the following table.
Medium Z [kg·m-2
·s-1
·106
] Medium Z [kg·m-2
·s-1
·106
]
air 0.0004 brain 1.55 – 1.66
water 1.52 blood 1.62
lung 0.26 soft tissues 1.6 – 1.74
fat 1.35 bone 3.75 – 7.38
Ultrasound needs for its propagation an elastic medium. Reflection and refraction can
occur when a beam of ultrasound reaches the interface of two media with different acoustical
impedances. At the interface, a portion of the ultrasound energy will be transmitted into the
second medium, and a part of the energy will be reflected. Assuming a zero angle of
incidence of the beam, the fraction R of the incident energy that is reflected is given by the
equation:
R = [(Z2 - Z1)/(Z2 + Z1)]2
, (2)
where Z1, Z2 are acoustic impedances of the two media.
Ultrasound will not travel an infinite distance through any medium without changes. A
part of the ultrasonic energy will be lost due to attenuation in the transmitting medium. The
attenuation of ultrasound in any specific medium depends on frequency. When frequency
increases, attenuation increases too. This dependence is linear in the tissue. The attenuation
generally includes absorption and scattering. The attenuation for a medium can be expressed
in dB/cm or by using the half-value thickness, the distance in which the initial intensity is
6
reduced by one half. The half-value thickness for bone at a frequency of 2.5 MHz is 6.5 mm.
For the brain, the liver and the kidney it varies between 15 - 30 mm.
At megahertz frequencies used in medical applications, ultrasound is generated and
detected by probes utilising the piezoelectric effect. The probes, also called transducers,
convert electrical signals into acoustical ones and vice versa. By applying an alternating
voltage of the desired frequency, the piezoelectric transducer generates an ultrasound signal of
that frequency. Conversely, when the transducer is periodically compressed by the reflected
ultrasound, an oscillating electrical potential arises according to the magnitude and the
direction of the deformation, and the same transducer becomes a receiver. The ultrasound is
emitted in short impulses.
The main physical mechanism of diagnostic application of ultrasound is its reflection at
acoustic boundaries in tissues. Ultrasound that is not reflected travels beyond the boundary
and may be reflected at deeper boundaries. The maximum penetration is limited by the
attenuation of the ultrasound passing through the given tissue. It is important to be aware that
ultrasound is almost completely reflected at boundaries with gas. This restricts the
investigation of gas-containing structures. For this reason, it is necessary to exclude air
between the probe and the patient. This is done by using a special gel or water or another
layer with closely similar acoustical impedance to the examined tissue.
Ultrasound diagnostics involves two groups of methods:
- Imaging methods
- Doppler methods
12.3.2 The mechanism of ultrasound imaging
All ultrasound diagnostic methods are based on the detection of ultrasound waves that are
reflected at boundaries between different tissues of the body. The magnitude of the reflected
wave depends on the difference of acoustic impedances of tissues forming such a boundary.
The very short ultrasound pulse (shorter than 1 s) generated by the piezoelectric transducer
and transmitted into the body (the mean repetition frequency is 1000 Hz) is reflected on the
boundary between two media of different acoustical impedance and detected by the same
probe. The time delay between the transmission of the pulse and the detection of its echo
depends on the propagation speed and the path length. The propagation speeds in different
soft tissues are so similar that a constant relationship between time and distance can usually
be assumed. The intensity of the echo compared to that of the incident pulse is strongly
diminished for two reasons:
1) The reflected part of ultrasound depends on the acoustical impedances of the tissues on
the boundary according to equation (2).
2) The intensity of the ultrasound decreases exponentially with the depth of penetration
according to the equation:
I = Io·e-·d
(3)
where Io is the initial intensity,  is an attenuation coefficient, d is the distance.
The attenuation due to the depth of penetration is compensated by special amplifiers. The
frequency dependence of attenuation limits the choice of scanning frequency. Lower
frequencies (2.5 - 6 MHz) are used for scanning of tissues located deep in the body, whereas
higher frequencies (7 - 15 MHz) serve for scanning of superficial organs.
Depending on echo processing, several scanning systems which may be used:
A-mode (Amplitude modulated): This basic pulse-echo system is used for measuring the
depth of echo-producing boundaries in one direction. The display on the monitor is called A-
7
scan and is characterised by deflections of the time base. The time delays between pulses
correspond to the distances between the boundaries within the body (Fig. 12.3.1a-a). This
one-dimensional imaging method is actually used for biometric measurements, especially in
ophthalmology.
B-mode (Brightness modulated): The information obtained by the transducer can be
processed and displayed on a brightness-modulated system so that the brightness increases or
decreases with the echo amplitude. It is the basis for two-dimensional imaging called B-scan.
The aim of B-scanning is the production of an image of a cross-section through the soft tissue
structures of the body (Fig. 12.3.1a-b).
In the older approach, the mechanical system restricts the motion of the probe to a twodimensional
plane and links the direction and the position of a B-scope time base on monitor
screen. This method is obsolete, the image on the monitor is not displayed in real time and
thus the tracing of moving structures was not possible (Fig.12.3.1b).
The current approach utilises a real-time scanner with multiple separate transducer
elements (up to 128). The elementary transducers are synchronously switched. By this
method, the need for mechanical motion of the probe is eliminated. The shape of the probe
exactly defines its ultrasonic field and simultaneously the shape of the cross-section of the
body that is momentarily displayed on the monitor. When the probe is moved over the surface
of the body, you can see images corresponding to the cross-sections of the structures situated
in the ultrasound field of the probe. Probes for linear and sector scanning are widely used
(Fig 12.3.1c). When the frequency is increased, the image resolution improves, but
penetration decreases. This problem is now solved by using multifrequency imaging probes.
These probes scan the near field with higher frequencies and the far field with lower
Fig. 12.3.1a. Principle of
A-mode (a) and B-mode
(b) imaging.
Fig. 12.3.1b Static B-scan
of epigastrium
8
frequencies. Probes differ according to the geometrical arrangement of elementary transducers
and their electric driving (phased array, linear array, annular array probes). Convex probes are
popular, the transducer arrangement on the convex area is almost linear, and the image,
however, is similar to that of a sector probe.
Sector and convex probes operating at lower frequencies are generally used to scanning
deep structures (abdominal scanning) – Fig. 12.3.1d. Linear probes operating at higher
frequencies are designed for examination of superficial organs.
M-mode recording (sometimes referred to as T-M mode). M-mode (motion-mode) is
used to examination of moving boundaries (heart valves are the typical example). M-mode
recording is composed of B-scans curves lying side by side. Every curve describes the motion
of one boundary expanded in time (i.e., one axis is time; second axis is the shift of the
boundary). Depending on the velocity of the movement and the depth of the examined
structure, it is necessary to choose an appropriate repetition rate. On the one hand, the deeper
the structure, the slower must be the repetition frequency; on the other hand, the faster the
movement of the structure, the higher must be the frequency of repetition. This implies that it
can be a problem to display fast and deep structures.
12.3.3 Doppler diagnostic methods
The measuring and imaging of moving structures is based on the Doppler principle. The
frequency shift between incident and reflected waves depends on the velocity of the
movement. If v is the velocity of the moving boundary, and c is the speed of propagation of
ultrasound in the given medium, the Doppler frequency shift fD in the reflected wave is given
by the equation:
cos
.2
c
fv
fD
= , (4)
Fig. 12.3.1c. Imaging lines of a
sector (a), convex (b) and
linear transducer (c)
Fig. 12.3.1.d. convex image of
liver region with gallbladder
and a small gallstone. An
acoustic shadow forms behind
the stone.
9
where  is the angle between the direction of the ultrasound beam and the direction of the
moving structure.
Ultrasound Doppler methods are now widely used in the study of moving structures or in
the measurement of the velocity of the blood flow in the vessels.
Continuous-wave Doppler (CWD) and pulsed-wave Doppler (PWD) systems, both
produce Doppler shift signals which depend on the velocity of the movement. The
continuous-wave system utilises separate transmitting and receiving transducers (Fig.
12.3.3a). This system can be used only for flow measurement in superficial vessels. The most
flowmeters are directional, e.g., they can measure separately direct and reverse flow (Fig
12.3.3b). The signal processed can be displayed in the form of the velocity time-course on a
monitor and simultaneously supplied to a loudspeaker.
The pulsed-wave system also provides information about the position of the examined
moving structure. This system is able to measure blood flow in deeply located vessels
independently on other moving structures in the surrounding. Recording of velocity curves is
often designated as spectral Doppler.
Most systems using the Doppler principle are combined with two-dimensional imaging.
This combination is called duplex method. The vessel under examination is first identified
using a two-dimensional scan and then the flow parameters are measured by the Doppler
method. In more sophisticated systems it is possible to co-ordinate the colours to the defined
range of velocities. This colour flow mapping (CFM) method enables us to improve the
analysis of the moving structures. Combination of a two-dimensional B-scan, spectral
Doppler and colour flow mapping represents a triplex method (Fig 12.3.3c,d). See also
chapter 12.3.7.
Fig. 12.3.3a. Velocity measurement
using CW Doppler
Fig. 12.3.3b. Velocity curves of a
directional flowmeter
10
12.3.4 Ultrasound echo-contrast agents
Ultrasound contrast agents (echo-contrast or echo-enhancing agents) are air or gas
microbubbles, free or encapsulated in a polymer cover. Echo enhancement made its first
appearance in 1968 when an increase in ultrasound signal intensity was observed after a rapid
intravenous injection of saline. However, the mechanism of this echogenic effect remained
obscure until 1980 when it was proved that echo enhancement is due to microscopic air
bubbles that scatter ultrasound energy in all directions. These bubbles have significantly lower
acoustic impedance than the surrounding liquid medium. This higher difference in acoustic
impedance can enhance the echogenicity of the body space in which they are introduced.
Short-lived bubble suspensions were first applied in echocardiography to enhance the
echogenicity of the right side of the heart. However, none of them survived pulmonary transit
to enter the left side of the heart and the systemic circulation. Therefore, efforts to produce
longer-lasting enhancement have been directed towards the creation of stable suspensions of
gas microbubbles that can pass through the pulmonary capillary bed. Three approaches have
been used to produce stable contrast agents:
- gas-filled microspheres
- microparticle suspensions
- suspensions of free gas bubbles in pure liquid vehicle
The factors that determine bubble survival are the rate of diffusion of the gas from bubbles
into the liquid, the viscosity of the fluid and their surface tension. Gas-filled microspheres are
protected against gas loss by a shell (for example of serum albumin or other biopolymer).
Water-soluble galactose microparticles originate gas microbubbles when they are mixed with
an aqueous vehicle – Fig. 12.3.4. Palmitic acid of 0.1 % forms a flexible protective monolayer
around the gas bubble and retards the diffusion of the gas into the surrounding medium.
The third approach is represented by a liquid in liquid dispersion in water that contains a
liquid, sulphur hexafluoride (SonoVue), which converts into gas microbubbles at body
temperature. The most gas bubbles in all types of contrast agents have a diameter of 2-5 m.
Fig. 12.3.3c. Duplex measurement of
blood flow in lower limb - a. profunda
femoris (high resistant curve)
Fig. 12.3.4. Electron-micrograph
(SEM) of an echo-contrast agent
Fig. 12.3.3d Triplex measurement of
blood flow in intrarenal artery (low
resistant curve)
11
Clinical applications example:
- Local application - hysterosalpingo-contrast ultrasonography. Intrauterine injection of a
contrast agent improves visualisation of the Fallopian tubes.
- Systematic use in Doppler ultrasonography. Intravenous injection of contrast agent
dramatically improves the amplitude of the Doppler signal. This application is particularly
useful in transcranial examinations, echocardiographic examinations, imaging of carotid
artery diseases and imaging of metastases of malignant tumours.
The improvement of image quality increases diagnostic confidence, leads to better
treatment decision, saves time, and reduces the need for invasive procedures.
12.3.5 Additional forms of ultrasound imaging
Harmonic imaging. About 15–20 % of patients seem poorly examinable by conventional
ultrasound imaging. To achieve a well evaluable image, it is necessary to increase acoustic
power of ultrasound impulses and the time of examination is also longer. A significant
increase of image quality in these patients and better contrast resolution in all other patients
can be achieved even without echo-contrast agents by means of a natural harmonic imaging
(THI, tissue harmonic imaging).
Principle of this method is very simple in first approximation: An ultrasound impulse with
fundamental (first harmonic) frequency f0 is transmitted in the body part. However, the probe
does not receive the reflections of the fundamental frequency but the second harmonic signals
with frequency 2f0. These oscillations are consequently processed. Unlike of the echo-contrast
harmonic imaging (when using the echo-contrast agents), the harmonic frequencies originate
directly in tissue structures as a result of so-called non-linear ultrasound propagation.
However, the technical realization of the measurement is not simple as the second harmonic
frequency is relatively weak.
Panoramic imaging. This imaging modality enables a continuous recording of the tissue
image across a broad area in desired direction. Hence an elongated view is produced for better
assessment of dimensions and morphology of the whole examined area (Fig. 12.3.5a). This
method is complementary to the conventional imaging which in most cases gives only partial
view on the examined body part.
Fig. 12.3.5a Panoramic image of epigastrium (from left: liver, right kidney, gall bladder, spleen)
Endoluminal imaging. It is a method of interventional radiology: miniaturised transducers
working at high frequency (20 - 40 MHz) are inserted into the blood vessels or hollow organs.
Their high-resolution ability allows detection of small pathological changes in the vessel
walls or some other organs. Fig. 12.3.5b.
12
Fig. 12.3.5b Endoluminal imaging: Transversal section of oesophagus. The numbers denote individual
layers of the oesophagus wall.
12.3.6 Ultrasound elastography
Elastography is a new imaging modality which is an emulation of the palpation. The basic
idea is that pathologic changes of tissues are manifested by their changed mechanical
properties, first of all rigidity. Malignant lesions are mostly more rigid in comparison with
benign lesion of healthy tissues.
The method provides images of tissue structures based on the measurement of the response
of the soft tissues on defined compression of the adjacent body surface. These properties
depend on molecular structure of individual tissue components (fat, collagen) and their spatial
organisation. Moreover, the tissues have also some viscoelastic and poroelastic properties
(poroelasticity is the specific elasticity of porous materials which pores are filled by a liquid).
It is not possible to obtain elastic properties of tissue from a simple ultrasound image.
Therefore, during last almost three decades several ultrasonic elastographic methods were
elaborated:
Static compression elastography or Strain-Stress Elastography, where the deformation of
the tissue is evoked by the pressure of examination probe (Fig. 12.3.6a).
Fig. 12.3.6a Static elastography of a phantom: a – scheme of the phantom, b- ultrasonogram, c-
elastogram.
13
In the Acoustic Radiation Forced Impulse Elastography the pressure is evoked by a strong
impulse of radiation force.
Transient elastography is a single-purpose method for diagnostics of the degree of fibrosis
in liver. Its typical technical component is a vibrator which produces mechanical middleamplitude
oscillations at a frequency of 50 Hz.
In present time, the SWE – Shear Wave Elastography has a dominant position in this field
of examination. Instead of pressure action of the probe, this method exploits the so-called
radiation force of the ultrasound waves. The acoustic compression is achieved by means of
relatively long repeated focused impulses along the imaging line, which produce acoustic
shear (transverse) waves. These waves propagate much slower than the longitudinal acoustic
waves and their speed is proportional to the elasticity of the tissue (Young modulus). The
particles of elastic medium move with amplitude of only micrometres. Imaging of such
movement needs a special imaging mode denoted as supersonic imaging with very fast image
processing (5000–20000 images per second). In contrary to the previous method the
information about tissue elasticity is quantitative and the colour scale is calibrated in kPa.
12.3.7 Special methods of Doppler imaging
Power Doppler (PD). The limitations of colour Doppler imaging are mostly removed in
the technology of the colour Doppler signal imaging denoted as "Power Doppler” or “Power
angio”. This method differs from the conventional blood flow imaging in utilisation of all
energy of the Doppler signal.
The advantages of this technology can be summarised as follows:
- Detection of the blood flow only little depends on the so-called Doppler angle and
enables us the flow imaging even at almost perpendicular incidence of the ultrasound beam on
the examined vessel.
Fig. 12.3.6b SWE of the
ductal carcinoma of the
breast. Stiffness 105 kPa
14
- There is no aliasing effect.
- The method can be used also for very slow flows hence it is “predetermined” for imaging
of organ of tissue blood perfusion (Fig. 12.3.7a).
Disadvantage: Missing information about the flow direction. Only one colour (mostly
orange) is used for flow coding. This disadvantage tries to remove a new method called
Directional Power Doppler.
Fig. 12.3.7a. Perfusion image of kidney obtained by Power Doppler. The bright areas are orange in
original picture.
B – flow. It is a new ultrasound diagnostic method based on combination of blood flow
and tissue structure imaging in real time. It is not Doppler method thus we obtain no
information about the blood flow speed. However, there is no worsening of spatial and time
resolution in comparison with conventional colour Doppler.
The B – flow method depicts blood flow similarly to static structures in grey scale. The
digitally encoded broad-band impulses reflect from the moving erythrocytes. The reflected
echoes are again decoded and filtered in order to increase the detection sensitivity. The main
advantage of this method is a correct imaging of the boundary between the blood vessel wall
and streaming blood. (Fig. 12.3.7b). In the conventional colour Doppler, the colour area often
overlaps the vessel walls. In superficially located vessels, e.g., carotids, this method can show
the position and extent of atheromatous plaques more precisely in comparison with colour
Doppler. Similarly, in the venous system we can see better small thrombi. The only limiting
factor of this method is the ultrasound attenuation which makes difficult imaging of deep
vessels.
Obr. 12.3.7.b Hepatic veins.
15
Tissue Doppler imaging. Originally, the movements of tissues, e.g., heart or vessel walls
or peristalsis of digestive tract could be observed and assessed only in a grey-scale image. The
Doppler imaging modality which is called Tissue Doppler Imaging (TDI), invented in 1994,
makes possible to get colour information about the speed and direction of tissue movement.
This method was invented for cardiologic purposes (detection of pathologic changes in
cardiac wall movement). However, today is used also in other areas of ultrasound diagnostics,
mainly in sonoangiology.
Basic principle: The ultrasound echoes reflected from moving tissues are relatively strong,
but the speed of the tissues is very small. On the other hand, the reflections from moving
erythrocytes are weak but the speed of blood is high. In the colour imaging of streaming
blood, the colour images of vessel walls and surrounding tissues are an interfering effect, a
colour artefact which is removed by filtration. In TDI are the signals caused by high speed of
streaming blood supressed, but the image is showing low speed movements (of heart or vessel
walls) in the range of 1 – 10 mm/s.
TDI given information mainly about diseases of coronary arteries, ventricular arrhythmias,
heart infarction etc. In angiology it enables more precise assessment of vessel wall elastic
properties connected with atherosclerosis. Other applications can be expected.
Advantages of colour duplex and triplex methods. The main advantage of these
methods is an easy and fast identification of a vessel in comparison with another part of the
body filled by a liquid. The tone colour, which brightness is a function of the streaming blood
velocity makes easier to find a stenosis and assess its degree. It makes also easier diagnostics
of pathologic changes of deeper vessels. In peripheral vessels, it makes the diagnosis more
precise, and in many cases can substitute X-ray angiography which endangers the patient by
ionising radiation.
A disadvantage of conventional colour mapping of blood flow is relatively small
sensitivity to slow streaming and a tendency to colour image artefacts which is caused by
additional movements (probe shifts, breathing or peristalsis movements) or by transmission of
arterial pulsation on the surrounding tissues. These colour artefacts can be removed or limited
by correct manipulation with the probe or by frequency filtering. Since these methods can
produce colour images of only mean stream flow it is possible that the maximal velocities can
be underestimated. Therefore, we must pay attention to the right setting of velocity range and
the record of velocity spectrum should be also added.
The general disadvantage of all colour methods is a relatively long-time interval necessary
for the production of the colour image (50 - 150 ms). It decreases the image frequency in
comparison with the grey scale image frequency. For example, if we need 65 ms to obtain a
colour image it means that the image frequency is only 15 frames per second. The quality of
colour image is also influenced by the size of the colour image window which is overlapping
the grey scale picture.
Many of these disadvantages or limitations can be removed by the new technologies of
image processing (e.g., power Doppler or B – flow method).
12.3.8 Principles of three-dimensional imaging (3D/4D)
The loss of one dimension, i.e., reduction of information about the volume element into a
planar two-dimensional image is a general disadvantage of all imaging methods. In
ultrasonography, an effort to remove this disadvantage by changing the image plane during
the examination appears recently. It can be done by means of probe movement during the
image acquisition. The probe can be shifted linearly, can be tilted, or rotated. The data about
reflectivity in individual planes are stored in memory of a computer which performs
16
mathematical reconstruction of the 3D-image based on sequence of planar images (Fig.
12.3.8a). Special probes for 3D imaging can be also used.
There are two basic algorithms for acquisition of 3D-image:
3D surface mode. It is based on identification of structure surfaces with sufficiently big
contrast to the surrounding tissues. The reconstruction is fast, but its information value is
relatively small. Imaging of foetus structural surfaces can be a good example (Fig. 12.3.8b)
Fig. 12.3.8a Reconstruction of 3D-image from a sequence of planar images.
3D volume mode is based on the vector spatial imaging of all the reflection points of the
examined structure which leads to formation of voxels. The position of each reflection point is
accurately specified in the image volume. An advantage of the volume image is a possibility
of its cross-section and rotation (Fig. 12.3.8c).
Fig. 12.3.8b 3D image of the foetus head
17
Fig. 12.3.8c 3D image of a tumour in oesophagus (dark area)
3D ultrasound multiplanar analysis could be the most accurate method for a correct
assessment of examined organs morphology. A digitalised 3D spatial image (Fig. 12.3.8d)
can be sent via suitable network to any chosen office or department.
Fig. 12.3.8d The multiplanar image of gall bladder region (from left to right: longitudinal,
transverse and coronal cross-sections plus volume reconstruction of the gall bladder)
The term 4D is used for systems which can process the volume image in real time. The
advantage of this kind of imaging is its complexness, the disadvantage is its relatively difficult
interpretation.
12.3.9 Safety of ultrasound diagnostic procedures
The interaction of ultrasound with biological tissues depends on its intensity. At
sufficiently high intensities so-called active interactions take place. These interactions are
used in physical therapy, surgery and in different laboratory experimental methods. Passive
interactions take place at low ultrasound intensities. These interactions form the basis of
ultrasound diagnostics.
Ultrasound of relatively low intensity has no ionising ability and its direct action on
biological system is effectuated by means of three main mechanisms:
- Heating due to the absorption of acoustic energy in tissues and its conversion into heat.
- Mechanical effects due to the radiation force and radiation torque in the vicinity of cell
and tissue interfaces.
- Cavitation meaning the production and dynamic behaviour of gas microbubbles formed
in a liquid medium during the underpressure phase of an acoustic wave. Bubble oscillation
and collapse cause microscopic shock waves capable of damaging surrounding cells and
forming free radicals.
From the practical point of view, the damage of tissues due to heating is more likely to
occur than that due to cavitation. For clinical practice, introduction of two on-screen indices
18
(i.e., numbers), the thermal (TI) and the mechanical (MI), is recommended. These indices
should warn the physician of potential risk.
12.3.10 Clinical value of ultrasonography
During the last sixty years, ultrasonography became the most widespread imaging
diagnostic tool in many branches of medicine. Ultrasonic examination forms mostly the first
step in the diagnostic imaging algorithm because of its accessibility and cost-effectiveness.
The basis of ultrasonography forms the dynamic (real time) B-mode imaging, enabling
detection of solid and/or cystic lesions in soft tissues. Only ophthalmology uses the onedimensional
A-mode imaging for precise biometric measurement. Very good results have
been achieved in applying ultrasound examination in cardiology, gastroenterology, obstetrics
and gynaecology and endocrinology. Musculoskeletal ultrasound is in the stage of rapid
development.
For assessment of haemodynamics and functional state of moving organs, sophisticated
duplex and triplex methods have been developed. These methods represent a combination of
dynamic B-mode imaging with spectral and colour Doppler methods. Several new methods,
such as intraoperative and interventional ultrasound, 3-D imaging, and harmonic imaging are
in the stage of technical development. These advancements should ensure that ultrasound
remains the preferred diagnostic imaging modality of the near future.
19
12.4 Endoscopic methods
Endoscopes comprise a group of optical devices used for viewing internal body cavities,
based on the reflection and refraction of optical beams. The endoscope is introduced in the
cavity under examination either by its natural orifice (nose and oral cavity, larynx, trachea and
bronchi, bladder, vagina, rectum) or by openings created surgically (thorax, abdominal cavity,
joint etc.).
Endoscopes can be classified according to their complexity, form of illumination and
mode of observation.
According to their complexity, endoscopes can be classified in three groups:
a) endoscopic mirrors,
b) endoscopes with rigid tubes,
c) fibroscopes and videoendoscopes, capsule endoscopy.
The source of light can be external or internal. Mirror endoscopes have external
illumination. In rigid tubes, the source of light forms an integral part of the device. The source
can be introduced within the tube in the cavity (distal illumination) or the light beam is
introduced in the cavity from outside (proximal illumination).
Observation of the cavity be can direct or indirect. In direct observation, the examiner
looks directly into the optical system of the device. In indirect observation, the image of the
cavity is scanned by a microcamera and displayed on the screen.
12.4.1 Endoscopic mirrors
Simple endoscopes are flat, convex, or specially shaped specular (mirroring) surfaces.
The curved surface focuses the light on the region of interest. The most frequently used
endoscopic mirrors are following:
The laryngoscope is a small flat mirror forming on a holder at an angle of 60°. It is
designed for examining the larynx and the (naso)pharynx.
The otoscope is a cone shaped mirror for visual examination of the external auditory
canal and the eardrum. It is equipped often by light source and magnifying glass.
The ophthalmoscope permits the physician to examine the interior of the eye. Bright
light is projected into the subject's eye, and the light reflected from the retina can be focused
by the examiner. The lens system of the patient's eye acts as a built-in magnifier (Fig. 12.4.1).
The vaginal speculum is a double spoon-shaped instrument with a mirroring part used in
gynaecology for examination of the vagina and the cervix uteri. Today it is a component of
the colposcope together with a camera and some mechanical parts.
Fig. 12.4.1. Principle of an
ophthalmoscope
20
12.4.2 Endoscopes with rigid tubes
These types of endoscopes have been developed prior to the discovery of fibre-optic
techniques (Adolf Kussmaul, 1868). They consist of metallic tube of different length with a
light source. Many of them are equipped with optical attachments to magnify the tissue being
studied (Fig. 12.4.2). Through the tube, special surgical instruments can be introduced for
taking tissue samples or other minor surgery.
Rigid tube endoscopes have two main disadvantages:
- The rigidity of the endoscope body represents a certain discomfort for the patient.
- The light transmission is connected with great losses of light energy.
Endoscopes often have names indicating their purpose:
Bronchoscopes for examining the air ways in the lungs.
Gastroscopes for examining the stomach.
Cystoscopes for examining the urinary bladder.
Rectoscopes (proctoscopes) for examining the rectum.
Laparoscopes for examining the abdominal cavity.
Arthroscopes for examining the articular cavities.
Many of these rigid endoscopes have been replaced by flexible ones.
Fig. 12.4.2. Longitudinal section of a rigid endoscope
12.4.3 Fibre-optic endoscopes and other endoscopic systems
The development of fibre-optic technique permitted the construction of flexible
endoscopes. These instruments can be used to obtain visual information from regions of the
body that cannot be examined with rigid endoscopes, such as the small and large intestine.
The light is conducted by optical fibre with minimal losses using the principle of the total
reflection.
An optical fibre is usually formed by two layers (Fig 12.4.3a). The core of the fibre has a
higher refraction index of n1 than the outward packing n2. Exploitation of the total reflection
depends on the angle α, under which the light beam enters the fibre. Total reflection takes
place if the following condition is fulfilled: sin < (n1
2
- n2
2
)1/2
, meaning that the angle of the
incident beam must be greater than the critical angle. Separate fibres for illumination and for
Fig. 12.4.3a. Schematic
picture of an optic fibre
21
image transmission form great bundles. In the image-transmitting bundle, the position of
single fibres on the input and output must be the same. Losses of light energy in optical fibres
are very low. The attenuation does not exceed 0.001 - 0.005 dB/m.
In the body of a fibre-optic endoscope, there are several channels: two optical channels
(light and viewing channels), water or air channel and a biopsy channel permitting passage of
devices to take tissue samples (Fig. 12.4.3b). On the distal end of the viewing channel is a
viewing objective. On the proximal end, there is a viewing ocular and a mechanical control of
the endoscope distal part. A halogen or LED source of light forms an integral part of the
endoscope.
Videoendoscope represents a more sophisticated version of fibre-optic endoscopes. In
this instrument the viewing objective is replaced by a micro-camera and the image of the
examined cavity is displayed on the TV screen (Fig. 12.4.3c).
The capsule endoscopes are small independent probes which can be swallowed. During
their passage through the digestive tract, they repeatedly record images and transmit them to
the external part which is connected with the monitor and storage unit. These devices are
disposable. Their independent part consists of a small battery, light source, camera, and
transmitter electronics.
Fig. 12.4.3b. Cross-section of a
fibre-optic endoscope
Fig. 12.4.3c.
Duodenovideoscope
(Olympus)
22
12.5 X-ray imaging methods
X-ray diagnostics began more than one hundred years ago with the epochal discovery of
German physicist Wilhelm Conrad Roentgen (in 1895), who obtained the first radiograph of a
part of human body, the famous picture of his wife's hand. In recent decades, the X-ray
examinations were partly replaced by ultrasound or radionuclide examinations, but X-ray
imaging nevertheless remains a commonly used diagnostic modality. Relatively low costs,
facility and precision are the advantages of the X-ray images. Exposing patients to ionising
radiation is the main disadvantage. Nevertheless, X-ray diagnostics still remain one of the
most important imaging diagnostic methods in most branches of medicine.
X-ray imaging methods are based on the principle of different absorption and scattering
of X-rays in different body tissues. Attenuation (i.e. absorption plus scattering) of a
transmitted beam of radiation can be expressed by the term
I = I0·e-·d
,
where I0 is the original intensity of radiation, I is the final intensity of transmitted radiation, d
is the thickness of the absorbing layer, and  is the attenuation coefficient. This coefficient
depends mainly on the effective proton number of the absorbing medium, which is given by
the mean proton number of elements present, considering their relative proportions. The
attenuation coefficient also depends on the energy of transmitted radiation, and the kind of
interaction of photons with matter (the photoelectric phenomenon, Compton scattering and
electron-positron pair production).
The so-called densitometry is based on the mere measurement of the attenuation of x- or
gamma-rays in bones. The method serves for determination of the degree of bone tissue
calcification. It is also called DEXA – dual X-ray absorptiometry.
12.5.1 Principal scheme of an X-ray device
In general, each X-ray diagnostic apparatus consists of several basic parts. Electrical parts
include a source of direct high voltage, X-ray tube, control panel and image detector. There
are also some mechanical parts, which change the position of the patient relatively the X-ray
tube and give mechanical support to the whole system.
The source of high voltage for feeding the X-ray tube (correctly, Coolidge type tube, see
below) consists mainly of a transformer, rectifier, and the circuit for smoothing the pulsating
direct current which is supplied by the rectifier. The transformer changes the relatively low
mains voltage (230 or 380 V) to the very high voltage necessary for feeding the X-ray tube
(roughly up to 150 kV). The output voltage of the transformer is continuously adjustable, and
its high values result in high values of the X-ray photon energy. This high voltage could be
led directly to the X-ray tube, but it is alternating, so that only each second half-period of the
alternating current could be utilised for production of the X-rays. Electrons can pass through
the X-ray tube in only one direction, i.e. from the hot filament (cathode) to the anode.
Therefore, the alternating current is changed into direct current by means of a rectifier.
However, the rectifier gives a pulsating current consisting of sinusoidal half-waves. The
changing voltage of these half-waves could cause unwanted production of photons of low
energy. That is why this pulsating current must be smoothed. The pulsation is almost fully
removed in this way.
Besides changing the voltage led to the X-ray tube, we can also change the intensity of
the current through the X-ray tube by changing of filament heating. The intensity of current
23
through the X-ray tube (the "amperage") influences the intensity of the X-ray beam, but not
the energy (i.e. the penetration ability) of individual photons.
The control panel (control unit) of the instrument is equipped by different control
elements and meters, mainly a voltmeter and ammeter to check the voltage and current across
the tube. There is also an ammeter for checking the heating current, time switches etc. Other
control elements serve for positioning of the patient and the instrument. In modern
instruments, most of these functions are controlled by means of a computer and its software.
In most cases, the control unit is placed outside the examination room, or behind a shield
made of lead glass, so that the useless and endangering exposure of the physicians and other
medical staff is avoided.
The main mechanical part of the instrument is a stand which holds the X-ray tube and its
shielding housing. The stand is usually mechanically connected with the examination table on
which the patient lies. The so-called Bucky grid (see Fig 12.5.2.1) is built in the table, along
with a plane of the image detector (see Fig. 12.5.3.1). If not necessary, the grid can be
removed. The table with the patient is moved manually or by means of servomechanisms.
While the electrical parts of X-ray instruments are fairly uniform, the mechanical parts
vary greatly depending on the use of the instrument. For example, movable instruments have
no examination table, but can be used directly in surgery or hospital rooms.
12.5.2 Origin of the X-ray image
X-rays are an electromagnetic radiation originating in atomic electron shells. They are
photons of very high energy (higher than the energy of ultraviolet light photons). They are
produced in X-ray tubes (Fig. 12.5.2), glassy evacuated tubes involving two electrodes: the
anode and the cathode. The construction of modern X-ray tubes differs from the Roentgen’s
original. The modern tubes are often called Coolidge tubes.
Electrons are intensely accelerated by a very high electric voltage in the space between
the hot cathode, also called the filament, and the anode. The energy of electrons incident
upon the anode is given by the term U·e. The electrons are suddenly decelerated in a tungsten
target which is a part of the anode. A small part of the liberated energy (kinetic energy of the
electrons decreases during deceleration) is transformed into high-energy photons - the X-rays.
Only seldom can happen that all the energy of electrons is transformed into energy of
photons.
,
where h·fmax is the maximum energy of the liberated photon (under the condition that all the
kinetic energy of the electron is transformed into photon energy). Such photons are of the
maximum possible frequency fmax or of the shortest possible wavelength min. The spectrum of
these X-rays is continuous. The X-rays arising in this way are also called “Bremsstrahlung“
(i.e. deceleration radiation in German). Photons of X-rays can also originate by jumps of
electrons between the innermost electron shells of heavy atoms, if there is a free place formed
by incident electrons. Their energetic spectrum consists of lines. This is the so-called
characteristic radiation.
24
12.5.2.1 Course of X-rays
X-rays are emitted from a very small area on the anode surface which is bombarded by
the electron beam coming from the cathode. This small area is called the focus. Thus, it is not
a point source of radiation. The X-rays propagate in straight lines in the surroundings of the
X-ray tube. The relatively narrow beam of radiation necessary for formation of the image on
the detector is delimited by the movable aperture cylinders and cones, usually made of lead.
The rays pass then through the body and hit the detector (or fluorescent screen or the
radiographic film). There they form electric signal or visible light image or a latent
photographic image as a consequence of unequal attenuation of the radiation beam by
different parts of the patient's body. The X-ray image is analogous to a shadow behind a
three-dimensional, semitransparent and optically non-homogeneous body placed in the path of
a light beam coming from an almost point source. The formation of an image depends on
different absorption (attenuation) coefficients and thicknesses of inner body structures. We
will describe the passage of X-rays through the body in more detail.
X-rays originating on the anode pass through the glass wall of the X-ray tube. Some
photons are absorbed there. Low-energy photons are then absorbed in the primary filter,
which is made of, e.g. aluminium plate of 3 mm in thickness. This filter absorbs mainly
photons which could not contribute to the image formation but could cause damage to the
patient's skin and subcutaneous tissues, where they are absorbed. The remaining, higher
energy X-rays are transmitted through the targeted body part, in which absorption and
scattering occurs. In many cases the so-called Bucky grid, (Fig. 12.5.2.1), which is near the
screen or film, is used.
The Bucky grid absorbs photons which were scattered in the body and no longer travel
parallel to the original bundle. The grid is made of parallel-layered lead strips, amongst which
material is nearly transparent to X-rays. The gaps between strips are so oriented that only rays
propagating in their original direction can pass through them. To avoid creating shadows of
the strips on the detector, they must be very fine, or they must move during the exposure. The
Bucky grid increases the patient's exposure about two to six times, but also absorbs 80 - 90 %
of the scattered X-rays.
Fig. 12.5.2. The X-ray
tube. K – cathode with hot
filament (main part of the
cathode, the whole
cathode is often denoted
as filament), W - tungsten
target.
Fig. 12.5.2.1. Principle of
the Bucky grid
25
The bundle of the X-rays finally comes to the place where the image is formed. If a
fluorescence screen is used for direct observation of the image, the imaging method is then
called fluoroscopy. Stable images, radiographs, usually took the form of a photo negative on
a developed photographic film. This method is called (general or plain film) radiography.
The blackening could be achieved directly by the X-rays, but by attaching fluorescent foils
(intensifying foils) to the film we can achieve a substantial enhancement of the blackening.
This way of image acquisition becomes obsolete. Modern digital technologies are based on
planar semiconductor detectors (usually called flat panel detectors). We can obtain both
fluoroscopic records and radiographic images using this type of detectors. There are two main
approaches of conversion of X-ray photons to electric signal in these flat panel detectors –
direct and indirect conversion. Direct conversion detector utilizes photoconductors, such as
amorphous selenium, to capture and convert incident x-ray photons directly into free electric
charge. Signal is then typically read out by a thin-film transistor array. This setting provides
high spatial resolution sensitivity for low energy photons, motivate the use of this detector
especially for mammography. Indirect detectors additionally contain a layer of scintillator
material, which converts the x-rays into visible light. The photoelectric converter is based on
amorphous silicon photodiode.
Flat panel detectors are directly connected to the X-ray device enabling automatic
optimalisation of exposition parameters. Direct digital radiography should not be confused
with computed radiography, which uses cassettes with photostimulable phosphor plates.
The plate is exposed to x-ray radiation exciting the phosphor, excited electrons are trapped in
the lattice until they are stimulated by the second round of illumination. Analog to digital
conversion is performed in separate machine, where the radiography plate is exposed to a
small, high-intensity laser resulting in the previously trapped electrons to return to their
respective valence bands, emitting light. This light is subsequently converted into an electric
signal.
12.5.2.2 Unsharpness of the image
Even a very good radiograph is not absolutely sharp. The boundaries between projections
of different inner structures of the body exhibit a continuous change of grey shades, even if
the boundaries themselves are well defined (for example, boundaries between the bone and
soft tissue). There are many reasons of this unsharpness.
(1) Movements of the patient. For example, tremor, movements by small children,
breathing movements, tissue shifts caused by pulse waves or heart action. This negative factor
can be limited by shortening the exposure time, but we need, of course, more intensive
radiation for that.
(2) The so-called geometric penumbra is caused by the non-zero area of the anode
focus. The source of radiation is not a point source. Therefore, the rays are incident onto the
boundaries of differently absorbing media under different angles, which causes unsharpness
(blurring) of their images. See Fig. 12.5.2.2.
The geometric penumbra can be limited by:
- Diminishing the focus area. This increases the need of anode surface cooling.
- Shortening the distance between the patient and the detector,
- Increasing the distance between the X-ray tube and the patient,
- Improvement of the Bucky grid function.
26
(3) Diffraction and scattering under small angles occur (e.g. the Rayleigh scattering).
Most scattered rays are absorbed by the Bucky grid, but not all of them, so that they can cause
the unsharpness.
(4) Light emitted by the luminescent layer on the surface of the detector illuminates not
only the nearest parts of the detector, but to a certain extent also the neighbouring parts (light
propagates in all directions).
12.5.2.3 Usage of contrast agents
Soft tissues exhibit relatively small differences in the magnitude of their attenuation
coefficients, and therefore are almost indistinguishable in the plain radiograph. For that
reason, contrast agents were introduced, mainly for visualisation of various cavities, blood
vessels, ducts etc. The contrast can be positive or negative, i.e., the attenuation of the X-rays
can be increased or lowered. Positive contrast can be achieved by substances containing
heavy atoms. For example, the suspension of barium sulphate (BaSO4, the atomic mass of
barium equals 137.33), denoted also as "barium meal", is used for imaging and functional
examination of the gastrointestinal tract. This substance can be administered per os or per
rectum. Substances containing iodine atoms are used for contrast imaging of blood vessels,
bile ducts and urinary system.
Some hollow inner organs can be examined by means of negative contrast, using air or
some other gases, namely carbon dioxide because of its high solubility in body fluids and fast
removal by breathing. The cavities are inflated by the gas, so that they can be depicted as
structures of very low X-ray absorption.
12.5.3 The most important methods in X-ray diagnostics
We can mention only some of the most important and commonly used radiographic
methods in this chapter. Radiography can be encountered in almost any clinical branch of
medicine. Very complex apparatuses are used, for example, in angiology (diagnostics of
blood vessels), cardiology or urology. The most modern instruments use digital processing to
achieve better resolution, contrast, contouring etc.
Digital subtraction angiography (DSA) is a classic example of these methods. The
principle of this method is simple but demands precise technology. Two radiographs of the
same region are recorded, which differ only by the presence or absence of a contrast agent. A
contrast agent in the blood highlights the respective blood vessels or the surrounding tissue if
there is internal bleeding. The two images are then digitally subtracted. In the resulting image,
we can see only the structures in which the two original images differ - an image of the blood
vessels, haematomas etc.
A specific method of X-ray diagnostics is the mammography which is extremely
important for screening of breast cancer. Elderly women should regularly visit the
examination centres. Besides the mechanical parts of the device which can compress the
examined body part, the method exploits also softer X-rays. (Acceleration voltage about 30
Fig. 12.5.2.2. The origin of geometric
penumbra - unsharpness of the Xray
image.
27
kV, molybdenum anode, often also molybdenum primary filter is used. The molybdenum
filter removes photons of higher energies hence the imaging is done by means of photons of
low energy around 20 keV.)
12.5.3.1 Image intensifier
Conventional fluoroscopy, i.e., the direct observation of the image formed on the fluorescent screen, was a
classical method of X-ray diagnostics. It was, of course, always connected with certain risk for the patients as
well as the medical staff because of large absorbed doses of X-rays. That is why image intensifiers were
introduced in practice. See Fig. 12.5.3.1.
Image intensifiers are large vacuum electronic tubes which contain the primary fluorescent screen, the
photocathode, anode, anodic (secondary) fluorescent screen and the electron optics. Photons of X-rays pass
through the patient's body and form a visible light image on the fluorescent screen. Emitted photons of visible
light cause the photoelectric phenomenon in the photocathode. The ejected electrons are accelerated by the
voltage across the photocathode and anode (15 - 22 kV) and directed by electron optics on the (secondary)
fluorescent screen. There they form a small, inverted, but about 10,000-times brighter image, compared to the
image of the primary fluorescent screen. This small picture can be observed by a camera and seen on a monitor.
The monitor is usually placed in another room. Thus, the medical staff is almost absolutely protected against
unwanted exposure. The considerable increase in image brightness enables us to also lower the patient's
absorbed dose to one tenth of the value of a conventional fluoroscopy.
Fig. 12.5.3.1. Scheme of the image intensifier. X - X-ray tube, O - imaged object, I1 - primary
image on fluorescent screen, G - glass support, FS - fluorescent screen, PC - photocathode, FE
- focusing electrodes (electron optics), A - anode, I2 - secondary image on the anodic screen, V
- camcorder. The individual parts are not proportionally depicted.
The image recorded by the camcorder can easily be digitised and stored in computer memory. Image
intensifiers are used by surgeons to introduce catheters or other tools and objects inside patients, and to check the
surgeon's work in orthopaedics, traumatology etc. In the late 1990s image intensifiers began being replaced with
flat panel detectors (FPDs) on fluoroscopy machines
12.5.3.2 X-ray apparatuses used in dentistry
Most X-ray apparatuses used in dentistry are very simple. They work with low-intensity
and relatively soft X-rays. They consist of a stand bearing the shielded X-ray tube and a
control unit. The X-ray tube must be placed close to the examined site, and the patients
themselves often use their fingers to keep a small piece of the film or a digital detector in
place. The digital detectors are connected by means of cables or wireless to a computer which
is able to process and display the image.
More sophisticated devices called orthopantomographs (OPG) give panoramic images
of maxillary or mandibular sets of teeth. The principle of imaging is very similar to layer
tomography.
In the last decade, an adapted cone CT system is also used in dentistry – see below.
12.5.3.3 Classical (layer) tomography
The classical or layer tomography was used relatively often in the past, but it is used only exceptionally
today, except of a very similar method in dentistry. It provides images of thin body layers, which were very
suitable for assessment of the position and shape of various lesions. In the simplest case, linear tomography, the
X-ray tube and the film cassette moved in parallel planes in opposite directions. Thus, the points lying in a plane
were projected on the same sites of the film. Only this single plane was sharply depicted. Structures lying outside
28
this plane were blurred. See Fig. 12.5.3.4.
The larger the angular displacements of the X-ray tube and the film (measured away the normal passing
through patient's body), the thinner was the sharply imaged layer (and the higher were the demands of the whole
system accuracy.
12.5.4 Computed tomography (CT)
Another X-ray imaging method - computed tomography (CT) - is one of the most
important imaging methods in medicine. The first patient was examined by this method in
London in 1971. The instrument was proposed by the English physicist G. N. Hounsfield, who
was awarded the Nobel Prize for medicine in 1979, together with the American physicist A.
M. Cormack who had proposed a similar principle (utilising gamma-rays) in the 1960's.
The principle of this method substantially differs from conventional X-ray imaging. The
image is not a "shadow" cast on a screen. It is a mathematical reconstruction by computer of a
transversal cross-section through the patient's body. CT-scanners of the first generation passed
a single narrow beam of X-rays through the body from the X-ray tube to a suitable detector
(scintillation or proportional counter) in the frame opposite (see Fig. 12.5.4).
The whole source-detector system moved linearly during the examination, so that the
beam passed gradually through a cross-section of the patient's body. After each such
movement (a scan), the system rotated by a small angle and a new scan was done. In the
second generation of CT-scanners, the narrow beam was made fan-like and able to hit an
array of detectors. The scanning movement and consequent rotation of the system was
maintained. In the third generation of CT-scanners, the detectors are arranged into an arc and
the whole system continuously rotates around the patient. Finally, in the fourth generation, the
detectors circle the patient. Only the X-ray tube moves. The further development can be
derived more likely from the third generation. Differences between the individual generations
of CT-scanners are shown in Fig. 12.5.4.
The time intervals much shorter than one second are today sufficient to obtain the picture
of one single cross-section. Despite of this fact, namely in connection with the 3D imaging,
the X-ray tubes used in the CT scanners must provide high power for relatively long time. It
results in strong heating of the anode which must have small focus, moreover. Therefore, it
must rotate and be quite massive. The imaging process can be accelerated (below 0.1 s) when
using two X-ray tubes, which beams are oriented in right angle. If these tubes work at
different voltages, it can be utilised for identification of materials of different composition,
e.g., to distinguish between the bone and the iodine contrast agent. This modification of CT
technology using two X-ray energy spectrums is called dual-energy CT. There are other
more possible ways how to construct dual-energy CT, e.g. rapid kV switching on X-ray tube,
dual layer of detector for “soft” and “hard” X-ray photons. Another evolution step in CT
material decomposition is photon-counting computed tomography. A photon-counting
Fig. 12.5.3.4. Scheme of the
classical layer tomography
29
detector registers the interactions of individual photons keeping track of the deposited energy
in each interaction.
The development towards 3D imaging took place mainly in nineties. It was conditioned
also by the progress in computer technology.
Spiral or helical CT scanners operate similarly to the systems of third or fourth
generation but there is additional shift of the patient’s body which enables the 3D
reconstruction of his or her body parts. We can imagine it as a layering of many 2D
tomograms hence we can speak about multiplanar imaging. Exposure time is up to tens of
seconds.
Similar results can be obtained by means of the devices equipped by detector systems
made of many parallel arcs. The number or arcs can by 128 or 254. In this way the exposure
time necessary for acquisition of a 3D image can be much shorter. The patient is irradiated by
an X-ray beam in the shape of a flattened cone to exploit all the arcs of detectors. This method
commonly called multi-slice CT or cone beam CT (because of the conical shape of the
beam. Some systems of this kind can work with lower resolution in the fluoroscopic mode
making so possible some interventions (catheterisation, punctures etc.). Of course, the doses
of radiation absorbed by patient’s body are considerably high and the time of the procedure is
limited.
A common problem of all CT systems providing 3D images are relatively high doses of
radiation, Therefore, the justification of this examination is very important.
It holds in all generations of CT-scanners, that a single X-ray beam passes through the
body, and its intensity decreases according to the mean attenuation coefficient of the passedthrough
tissues. The absorption profiles are recorded during the source-detector movement,
or simultaneously from the detector array. The detector data (absorption profiles) are digitised
and mathematically processed by the computer and displayed as a cross-sectional image,
which is actually the "map" of the X-ray attenuation at each point of the cross-section. This
map actually consist of matrix of voxels (volume matrix elements), usually with 512×512
definition. Computer processing reveals half-per-cent differences in attenuation values, which
is impossible in the conventional X-ray imaging. Soft tissues and their pathological changes
can be examined very well. Large demands on computer equipment are made by the threedimensional
image reconstruction in spiral CT-scanners.
The image (scan) on a computer screen is composed of points of different grey shades
which represent tissue attenuation coefficients converted into Hounsfield units (HU):
,
Fig. 12.5.4. The four generations of
CT-scanners.
30
where T is the attenuation coefficient of a medium (tissue), W is the attenuation coefficient
of water, and k is the constant of numerical value of 1000. It follows from this definition that
water has a value of zero in the HU scale and air about -1000, because the attenuation of Xrays
in air is negligible. The HU value for compact bone is about +1000. Therefore, we have a
range of 2000 HU available. It makes no sense to represent this whole scale by individual
grey shades because the human eye is able to distinguish only about 250 grey shades. Virtual
colours can be used, but it is usually not necessary, as most tissues exhibit HU values from
zero to +100. In such an absorption window, a grey scale is sufficient (see Table 12.5.4).
It is very important that the HU values do not depend on the energy of X-photons
(given by acceleration voltage) so we can compare images obtained in different CT systems.
Very good resolution of soft tissues, including tumours, is the main advantage of CT
imaging. The method also is very convenient for planning surgery and radiotherapy of
tumours. The resolution can be improved by means of the contrast agents. The disadvantage is
the approx. ten-time higher absorbed dose of radiation in comparison with conventional
radiography. CT-scanners are also expensive, and they must be operated by highly trained and
specialised staff.
Tissue or material Attenuation
coefficient  [cm-1
]
HU
water 0.22 0
fresh blood 0.23 47
blood coagula 0.24 62
erythrocytes 0.24 76
blood plasma 0.22 11
white brain matter 0.23 30
grey brain matter 0.23 34
white matter oedema 0.23 25
bone (cranial) 0.38 706
neutral fat 0.21 -70
cerebrospinal fluid 0.22 5
Table 12.5.4. Attenuation coefficients and values of Hounsfield units for selected tissues.
31
12.6 Radionuclide imaging and other diagnostic methods
Radioactive elements, radionuclides, can be used in medicine in many ways. This chapter
will deal with only some of them which need special instrumental technology. In some cases,
the simple detectors (see section 11.8) are sufficient.
We can define the following main regions of radionuclide utilisation in biomedical
sciences:
- tracing
- radioimmunoassay
- examination of physiology of body organs
- imaging of body organs or parts
In principle, only physiological examinations and imaging methods need special
instruments (gamma-cameras, the PET- and SPECT-scanners).
12.6.1 Tracing and radioimmunoassay
Radionuclide tracing is often used to calculate compartment volumes, such as the volume
of body water, circulating blood, or fat tissue. A known amount (known activity) of the
radionuclide is introduced into the organism, either orally or by injection. Then sufficient time
is allowed for the radionuclide to distribute evenly (i.e., reach equilibrium concentration) in
the whole compartment. By taking a small sample volume and measuring its radioactivity, it
is possible to calculate the total compartment volume. A similar condition must be also
fulfilled in some organ examinations and imaging methods. It is also possible to start with
activity measurement in the respective organ immediately after administration of the
radionuclide in the so-called dynamic examination. The diagnostic value is different, of
course.
Radioimmunoassay (RIA) is used in clinical biochemistry and haematology. It is used
for determination of trace amounts of substances, notably hormones, in blood samples taken
from patients. An interaction of antigen-antibody type is studied in vitro. The antibody is
labelled by a radionuclide.
In RIA and tracing, beta-emitters are mainly used (tritium, iodine-125, iron-59 etc.),
because the detector can be positioned close to the radioactive sample (we do not need
penetrating radiation).
12.6.2 Scintillation counters
The scintillation counter is, in principle, a relatively simple instrument consisting of a
scintillation detector, i.e. the counter proper (see Section 11.8.3), the mechanical part and the
lead collimator which is connected with the lead shielding of the detector. The collimator
makes it possible to detect radiation only from certain, relatively narrow, and sharply
delimited solid (spatial) angle. Therefore, other radioactive sources present in the body or in
its surroundings cannot influence the measurement results. Voltage pulses generated by the
scintillation detector represent individual photons of gamma-radiation. They are amplified,
counted, and recorded by joined computer. The examination by means of simple scintillation
counters can be performed in two ways, in principle:
(1) The detector is stable. It measures the rate of storage or capture of a radionuclide
bound to a suitable metabolite in a given place. Thus, this kind of examination gives
information about the time-course of local metabolic activity.
(2) The detector moves in lines or scans ("zigzags") above the patient's body part. Such
an instrument was denoted as a rectilinear scanner. So a "map" of the radionuclide
32
distribution could be obtained. At present, rectilinear scanners are replaced by more accurate
imaging radionuclide methods.
12.6.3 Anger gamma-camera
The gamma-camera, invented by Anger in 1958, is a special kind of scintillation detector.
The scintillation occurs in a large scintillator, a crystal of sodium iodide in the form of a
square or a disc up to 40 cm in diameter. Numerous photomultipliers or other detectors (up to
50) are placed on one side of the scintillator. Detected signal is digitised and processed by a
computer.
The positions of scintillation events, which reflect the distribution of a radionuclide inside
the body, can be determined by means of computer processing of the signal coming from
photomultipliers. However, a very important condition must be fulfilled to obtain an image: a
defined point in the scintillator must represent a defined point in the body. (Let us imagine the
plane of the scintillator as parallel with the plane in which the scintillator "sees" the projection
of a body part. The defined points lie in these two planes.) Such condition can be fulfilled
only using collimators. See Fig 12.6.3. So-called "pin-hole" (radiation passes through a small
hole in conical lead shield placed in front of the scintillator) or absorption collimators are
most often used. Absorption collimators are made of lead lamellas similarly to the Bucky grid
in conventional radiography.
Anger cameras are very effective diagnostic tools because they show the distribution of a
radionuclide inside the body very quickly. Therefore, they can be used also for imaging of
relatively fast processes, e.g., the passage of blood through coronary arteries. These cameras
can also move along the body axis, so we can quickly obtain an image of the radionuclide
distribution in the whole body. Metabolic pathways can be studied in this way. It is also
possible to search for cancer metastases if they are able to capture a substance containing a
gamma-emitter: today technetium-99m or iodine-123. The technetium radionuclide exhibits a
short half-life (6 hours in comparison with 8 days in the originally used iodine-131), so it
must be prepared directly in departments of nuclear medicine in technetium generators.
These generators are containers with radioactive molybdenum which decays to form
technetium-99m, which is obtained from the container by a chemical process. Both
radionuclides give radiation of relatively low energy, which is not able to evoke secondary
radioactivity.
In the examination of the thyroid gland, radioactive iodine-123 (its half-life is 13,22
hours) is used in the form of sodium iodide. The kidneys are examined by the technetium-
99m labelled compounds. Technetium-99m is not easily available, but it is almost an ideal
radionuclide because of its fast excretion from the body, the short half-life and emission of
almost pure gamma-radiation. The product of the Tc-99m decay is also little radioactive.
12.6.4 SPECT and PET
SPECT (single photon emission computed tomography) and PET (positron emission
Fig. 12.6.3. Schematic drawing of
Anger cameras with different kinds of
collimators. E - emitters, P photomultipliers,
S - scintillation
events (flashes), L - lead shielding of
the collimator.
33
tomography) are the most advanced diagnostic methods which utilise radionuclides. In
computed (X-ray) tomography, the source of radiation is outside the body. In SPECT and
PET, it is directly inside the body, but the way of image reconstruction is analogous.
In SPECT a source of gamma-radiation, i.e., the radionuclide, is introduced into the body
in a labelled compound, which can be captured and stored in some organs, or quickly
eliminated from the body (mainly with urine). Individual photons are counted by multiple
scintillation counters.
The counting is done stepwise from different directions, which allows the twodimensional
reconstruction of a body cross-section. The structures seen in this cross-section
are points or small volumes emitting the radiation.
The following detector arrangements and kinds of their motion are or were utilised for
SPECT:
a) A simple scintillation detector equipped by a collimator scanned and circled around the
body, similarly to the X-ray tube and the detector in the first generation of CT scanners. This
method is obsolete.
b) More scintillation detectors are arranged into a circle or a square around the body. The
whole system can rotate around the patient and move along his or her body axis. Also this
method became obsolete.
c) The Anger camera with one, two or three “heads” circles around the examined patient.
This way of image acquisition in used today.
SPECT uses the common sources of gamma-radiation (iodine-123 and technetium-99m)
but PET must use positron emitters. These are radionuclides of biologically most important
elements produced by bombardment of atomic nuclei by high-energy beams of accelerated
particles. The half-lives of these radionuclides are very short, 109,7 min. in fluorine-18, 20
min. in the carbon-11, or even shorter in some other radionuclides. The construction and
operation of the accelerators are very expensive. Because of the relatively long half-life of F-
18, it became the most often used radioisotope which can be transported by a car up to about
150 km from the producer’s accelerator (cyclotron).
The scintillation detectors are arranged in circles around the patient without the need of a
lead collimator. The band of detector circles is about 25 cm broad in the cranio-caudal
direction. Positrons can travel only very short paths in a liquid or solid (dense) medium
because they annihilate immediately when colliding with an electron, forming two photons of
gamma-radiation (E = 0.51 MeV) – see Section 2.1.3.3. These photons propagate in opposite
directions from the place of their creation. Such an annihilation event can be detected by two
detectors in so-called coincidence circuit, i.e., by detectors on opposite sides of the place of
annihilation. The impulse is counted and processed only if the photons are detected by an
opposed pair of detectors at the same time. (Photons propagate at the speed of light, so we can
neglect small differences in distance - annihilation always occurs closer to the first or second
detector). The principle of PET can be seen in Fig. 12.6.4a.
The resolving power of PET can be twice higher than SPECT.
The positron emitters can be bound in the metabolites, e.g., glucose derivatives, so it is
possible to obtain not morphologic but functional information. PET-scans of the oncological
patients are most frequent. However, we can also reveal brain centres which are actually
active, i.e. have an increased uptake of a glucose derivative. Along with the EEG, PET is one
of few methods by which it is possible to follow objectively the superior nervous activity of
brain centres.
34
Fig. 12.6.4a. The basic principle of the PET. The opposing detectors are
connected into a coincidental circuit. The radiation source S can be detected
only when lying on the connecting line of the two detectors. The detector A
can be hit from the source S2, but not the detector B, because the source S2
lies outside its detection angle. In SPECT, the signal coming from source S2
would be overlapped by the signal coming from the source S1. It explains the
high resolving power of PET.
12.7 Magnetic resonance imaging
12.7.1 Nuclear magnetic resonance
Magnetic resonance imaging (MRI) which is based on the spatial analysis of the
phenomenon of nuclear magnetic resonance (NMR) in human organism is probably the most
advanced diagnostic imaging method in present time. The first MRI scan (of a human chest)
was obtained by Damadian in 1977, but many other scientists took a part in the construction
of first devices during seventies. The physico-mathematical description of this phenomenon
and method is very complex and cannot be introduced here without unacceptable
simplification. This description will be mostly qualitative. The approaches of classical and
quantum mechanics will be also combined.
Each atomic nucleus with an odd number of nucleons (or even number of nucleons at an
odd number of protons) exhibits a certain magnetic moment , which can be imagined as a
consequence of the rotation of electrical charges of elementary particles forming the atom
nucleus - their spin. (In classical physics, the magnetic moment is defined as the product of
current intensity in a conductor loop, and the area delimited by this loop. It is a vector
perpendicular to the plane of the loop. It determines the torque acting on the loop in the
magnetic field of magnetic flux density B.) In practice, the light atomic nuclei with an odd
number of nucleons are the most interesting, with spin values of ± 1/2. The proton (atomic
nucleus of hydrogen), phosphorus P-31, carbon C-13, fluorine F-19 and sodium Na-23 are of
considerable medical importance. It should be noted that there is a direct proportionality
between the magnetic moment of nucleus  and its angular momentum S:
 = S,
where  is the so-called gyromagnetic ratio. The magnetic moment of nucleus as well as the
angular momentum of nucleus is vector with a direction identical with the axis of nuclear
rotation (nuclear spin). In the absence of an external magnetic field, the nuclear magnetic
moments are randomly oriented. Their vector sum per unit volume - the vector of
magnetisation - is zero.
Let us place the nuclei exhibiting a non-zero magnetic moment into a strong external
stationary and homogeneous magnetic field B. The nuclear magnetic moments will "try" to
35
orient in (or against) the direction of vector B. It is useful to identify this direction with the zaxis.
This orientation change will evoke a torque which will be demonstrated by the so-called
precession of nuclei - see Figs. 12.7.1a and 12.7.1b. The axis of nuclear rotation, more
precisely the vector of their angular momentum and hence the vector of magnetic moment,
will start to move like a flywheel (gyroscope), with its axis deflected from its original
direction by an external force. The vectors of magnetic moments of nuclei will start to
circumscribe the surface of a cone with its axis directed parallel to the vector B. The
frequency f of this precession motion, the Larmor precession, is denoted as the Larmor
frequency, and given by the formula
f = B/2π.
The nuclear magnetic moment is the "quantised" quantity. For nuclei with typically an
odd number of nucleons placed in an external magnetic field, only two energetic quantum
states are possible. The state with the spin value of +1/2 (, more precisely its projection into
the z-axis, and vector B are of the same orientation) is of slightly lower energy than the state
with the spin value -1/2 (z-projection of  and vector B are of the opposite orientation).
Therefore, in a given assembly of nuclei, the states characterised by spin value +1/2 are
slightly prevalent after reaching thermal equilibrium. The vector of magnetisation (its zcomponent)
will not equal zero in this case.
The precessing nuclei can absorb a quantum of electromagnetic radiation (a photon) of
frequency equal to the Larmor precession, and jump into a higher energetic state. That is
why the application of a radiofrequency pulse of such a frequency results in an increase of the
number of higher energy states in the whole "population". The vector of magnetisation will
not equal zero. Its z-component (the longitudinal magnetisation) will be oriented in opposite
direction. The precession motion will also be simultaneously harmonised in phase, and in the
Fig. 12.7.1a. States of atomic
nuclei with an odd number of
nucleons before and after
placement in a magnetic field.
Fig. 12.7.1b. Precession of resonating nuclei. P - hydrogen nucleus
(proton), B - magnetic field vector (density of magnetic flux), z - axis
identified with the axis of precession (parallel with B),  - magnetic
moment of nucleus, L - projection of the nuclear magnetic moment
in the direction of z-axis (their vector sum in unit volume is called
"vector of longitudinal magnetisation"), T - projection of the
magnetic moment into the plane xy (their vector sum in unit volume
is called "vector of transversal magnetisation").
36
xy-plane rotating component of magnetisation (the transversal magnetisation) will appear. We
could imagine that the vector of magnetisation will be also brought in precession motion. The
return to the low-energy ground state (relaxation) is possible in two ways: by relaxation which
is not connected with emission of electromagnetic radiation, and by emission of quanta of
electromagnetic energy, the NMR signal.
We can speak about two relaxation times. The first of them, the longitudinal relaxation
time T1 is the time necessary for the return of nuclear population into the original state with a
slight prevalence of the direction of longitudinal magnetisation identical with the orientation
of vector B. It is not the time of full return, which is measurable only with difficulty, but the
return of 63 % of the nuclei to the original ground state. This time is strongly influenced by
the interaction of magnetic moments with fluctuating magnetic fields of neighbour nuclei, so
that we speak about spin-lattice relaxation. T1 has values ranging from 300 to 2000 ms in
biological media. The transversal relaxation time T2 is two- to ten-times shorter than the
time T1.
During the transition of the nuclear population into the higher energy state, the
harmonisation in phase (coherence) is manifested by the non-zero value of transversal
magnetisation. The transversal or spin-spin relaxation time (T2) is the time necessary for
"dephasing" of precession and the renewal of the original zero value of transversal
magnetisation. Correctly, it is the time necessary for the decrease of transversal magnetisation
to 37 % of the maximal value achieved.
The relaxation times represent information additional which can be obtained from the
time-course (decay) of the NMR signal just after the radiofrequency pulse, and also from the
MRI scans described below. The intensity and also the repetition frequency of applied
radiofrequency pulses, in relation to the values of relaxation times, determine whether the
NMR signal will carry information about the spatial distribution of protons or about the time
T1 or T2. More detailed analysis of these problems is out of scope of this text. The presence of
paramagnetic atoms, for example gadolinium, considerably reduces the relaxation time T1,
which causes amplification of the signal. This is the basis for the utilisation of gadolinium and
some metals as contrast agents for MRI. Gadolinium compounds differ according to the
body part examined.
The Larmor precession of nuclei in magnetic field can be measurably shifted by their
chemical surrounding. This chemical shift is expressed in parts per million (ppm) of the value
of a chemical standard. For example, there is a considerable difference in chemical shifts of
protons bound in =CH- or -CH2- groups, and the effect of their more distant surroundings is
measurable. The measurement of chemical shifts is an important tool of structural analysis in
chemistry, and, as shown further, it is also of medical interest. The method of chemical
analysis based on the measurement of chemical shift is called NMR-spectroscopy.
12.7.2 The principle of image formation
The phenomenon of nuclear magnetic resonance can be induced and measured, in
principle, in two ways. In the first case, we can have an external magnetic field of constant B
and search for the energy (frequency) of electromagnetic oscillations able to cause nuclear
resonance. In the second case, which is easier to achieve we can work with a constant energy
(frequency) of electromagnetic oscillations and search for a value of B, at which resonance
occurs. In this case we change the Larmor precession frequency. This second possibility is
also exploited in magnetic resonance imaging – MRI.
When placing the examined body part in the homogeneous magnetic field, the
radiofrequency pulse will produce an NMR-signal in the whole of the examined part, and the
information about local characteristics of resonance will be lost. The situation is different in a
37
non-homogeneous magnetic field. When creating, for example, a field gradient in the
direction of z-axis (identified with the body axis, in practice), the resonance condition will be
fulfilled only in a thin slice of the examined part lying in the xy-plane - the signal will be
produced only by nuclei present in this slice. We can also create a field gradient in the
direction of x- or y-axis, forming so in the slice a thin stripe parallel with the x- or y-axis, or
identifying a point, respectively - see also Fig 12.7.2. The setting of these gradients is done in
various pulse regimens, and it is different in different variants of MRI or instruments.
However, in each case, it is possible to obtain spatially specific information about the
magnitude of a resonance signal, which is among other things proportional to the number of
resonating nuclei in a given part of space. Two- or three-dimensional information about the
distribution of resonating nuclei can be obtained by means of algorithms used in CT-method,
or based on the Fourier transform.
Not only differences in the amplitude of resonance signal, but also differences in
relaxation times can be depicted. We can also determine the chemical shift of resonating
nuclei in chosen parts of the image.
Fig. 12.7.2. The function of magnetic field gradients in image acquisition. A)
the plane of section (the slice) is obtained by application of the magnetic field
gradient, i.e. the setting of resonance value of B (and the Larmor frequency) in
the direction of z-axis, i.e. the patient's body axis. B) We can delimit a column
(stripe) by the gradient applied in the direction of x-axis in this slice. C) The
same done in direction of y-axis allows resolving another stripe.
Let us note some technical aspects of MRI instruments. It is necessary to work with
magnetic fields B which range from 0.1 to 3.0 tesla (T) to obtain the MRI scans. Devices
equipped by giant permanent magnets weighing even tens of tons can achieve B values up to
0.3 T. These devices were relatively cheap and had low operational costs, but their resolution
power was also relatively low, and they are not used more. Devices equipped by
electromagnets can reach rather higher B values and also better resolution, but the winding of
electromagnets must be cooled, and they consume a lot of expensive electric energy. Devices
with superconducting magnets have the best resolution, but they need liquid helium and hence
their operational costs are very high. These devices allow us to perform chemical analyses on
the basis of chemical shift determination.
The necessary gradients of magnetic field are achieved by means of additional (gradient)
coils, which disturb the homogeneity of the external magnetic field (its B is about 1 T, while
the typical values of B gradients are several mT per meter). Near the patient are transmitting
and receiving coils, i.e. the radiofrequency pulse source and receiver of NMR-signal of the
same frequency (tens of MHz).
MRI devices generate strong magnetic fields and disturbing electromagnetic signals. That
is why the coils and the electromagnet windings are maximally shielded. Despite this
shielding, the gantry (tunnel) of the device is a source of such a strong field that ferromagnetic
objects (e.g. small medical tools) can be pulled inside the gantry at very high speed and injure
the patient or damage the device. No microelectronics or vacuum electronics (CRT screens),
38
magnetic memory media, including magnetic paying cards, etc., can be placed within many
metres around the device. The shielding requirements are given by hygienic norms for these
devices. Conversely, the examination can be disturbed by external magnetic or
electromagnetic fields. Therefore, the examination rooms must be shielded (Faraday cage).
12.7.3 Clinical value of MRI
The MRI scan, i.e. the depiction of the tissue section, is a spatial reconstruction of
resonating nuclei density. In contrast to computed tomography, these sections need not be
only transversal, but can also be sagittal or frontal. The method gives very high contrast
resolution of soft tissues (with maximum resolution in order of tenths of millimetre) due to the
different densities of protons (hydrogen nuclei) in various body tissues. See Fig. 12.7.3.
In contrast to other tomographic methods, MRI does not utilise the ionising radiation,
which is a large advantage. No biological effects of applied strong magnetic fields or
radiofrequency pulses have been observed up to now, at least at the application times used in
standard MRI examination. For this reason it is possible to use MRI in children, and if
necessary in pregnant women, except during the first trimester of pregnancy.
There are, however, some potential problems for patients. They can be disturbed by the
considerable noise which accompanies the examination. Of course, no ferromagnetic
materials (old steel implants, iron splinters from war injuries etc.) or electronics (pacemakers
with no guarantee that they can be used in MRI under certain circumstances) can be present in
the bodies of examined patients. Problems can also be caused by some cosmetic preparations
or jewellery. In some devices, and in sensitive patients, the fear of closed space
(claustrophobia) can appear. However, it can be moderated by sedatives.
Some parameters of NMR-signal depend on temperature, for example, relaxation times or
the chemical shift of protons present in water molecules. (It depends on the state of hydrogen
bridges which are influenced by temperature.) It is possible to create an examination
algorithm which leads to the tomographic depiction of temperature values distribution. A
unique tool is obtained in this way, which enables us to measure directly and non-invasively
the changes of tissue temperature during, e.g. ultrasonic or laser tissue heating.
The NMR-signal is also sensitive to the movement of resonating nuclei. For example, in a
given site the nuclei present in streaming blood and "prepared" are always replaced by
"unprepared" nuclei, i.e. nuclei which do not have the required value of Larmor frequency.
We speak about a "washout effect" in this case. The analysis of signals coming from
Fig. 12.7.3. „T2 weighted“ image
of transversal section of head in
the level of cochlea. (Siemens)
39
structures in motion allows us to measure their velocity, e.g. the velocity of the blood stream
(magnetic resonance angiography).
The movement of depicted structures also causes, however, some unwanted artefacts, so
that the synchronisation with breathing or heart rate is necessary in some MRI examinations.
MRI can also be used for the study of metabolism, e.g. the metabolism of ATP. ATP
contains phosphorus P. The phosphorus chemical shift varies during the hydrolysis of ATP to
ADP + phosphate. The chemical shift of the free phosphate is also different. This information
can be obtained from individual small volumes of the examined body part, and the metabolic
activity, during which the ratio of ATP/ADP changes can be so visualised.
We already mentioned (Section 12.7.1) the use of gadolinium as an MRI contrast agent.
Gaseous microbubbles used as a contrast agent in ultrasonography are also a suitable MRI
contrast agent. (The pressure influences the bubble size and consequently medium density.
Therefore, the NMR signal depends considerably on the ambient pressure.) In future, it may
be possible to exploit these microbubbles for non-invasive MRI determination of actual blood
pressure anywhere in circulation.
As mentioned, MRI is the most advanced diagnostic method available. The development
of this technology was not finished. We may see the first real-time examinations, analogous to
dynamic ultrasonography, already.
However, the very high price and high operational costs limit the use of MRI. That is why
other imaging methods will remain in use for a very long time, e.g. conventional X-ray
diagnostics and CT mainly in examination of tissues with low content of water; in cases, in
which the MRI is contraindicated, ultrasonography, for its operative use in real-time and its
relatively low costs; and radionuclide methods, for their considerable ability to follow the fate
of metabolites in our bodies. Each of these methods also gives a different kind of information.