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Thracian University - Stara Zagora

MEDICAL FACULTY

Department of Physics, Biophysics, Reontgenology and Radiology

Facultative course

NEW TECHNOLOGIES IN MEDICINE

Content and Questionnaire

1. Physical basis of ultrasonography. 2D-echography. 2. Physical basis of Dopplerography imaging. 3. Application of lasers in medicine. 4. Linear accelerators for radiotherapy of tumors. 5. Physical basics of magnetic resonance imaging (MRI). 6. Radionuclide methods of diagnosis: positron-emission tomography (PET).

All questions from 1 to 6 should be read carefully and with understanding. Select only one question and learn it well. During the exam, the selected question has to be represented in written form and explained orally in sufficient details. 1. PHYSICAL BASIS OF ULTRASONOGRAPHY. 2D-ECHOGRAPHY

Ultrasound is a mechanical wave (Fig. 1), which is used in medical diagnostics for imaging of internal organs. The main advantage of ultrasonography consists in the considerably smaller hazards for human tissues compared to the diagnosis based on the X-rays and other ionizing radiations.

Fig. 1. Cycles of compression and rarefaction of the density and pressure in the medium caused by the propagation of continuous-wave ultrasound.

Continuous-wave (CW) ultrasound is described by the following parameters: frequency (f), period, wavelength (λ), propagation speed (C), amplitude and intensity. The higher the frequency, the shorter will be the wavelength and the smaller will be the size of the objects that we can diagnose. An important parameter of the propagation medium is the acoustic impedance, Z = ρ × C, where ρ is the density of the medium. Impedance is the relationship between acoustic pressure and the speed of particle vibration. The intensity (W/cm2) of ultrasound is the ratio of the acoustic power of the wave to the area of the surface through which the ultrasound flows.

Reflection, refraction, scattering and absorption of ultrasound

In ultrasound diagnostics the ultrasound passes from one tissue to another. Provided the adjacent tissues have different acoustic resistances and different speeds of conducting ultrasound, several phenomena will take place at their boundary surface including reflection, refraction and scattering.

Fig. 2. Reflection and refraction of ultrasound.

Intensity reflection coefficient is the ratio of the intensities of reflected and incident wave. The intensity of the ultrasound wave, transmitted through the interface, to that of the incident wave is the intensity transmission coefficient. These coefficients depend strongly on the difference in the acoustic impedances of both tissues. The ultrasound will not be reflected if the tissues have equal acoustic impedances, although their densities are different. On the other hand, when the difference between the acoustic impedances of both tissues is too large the intensity of the reflected wave tends to equal that of the incident one. An example of this is the air/skin boundary, where almost total reflection of the ultrasound takes place. Hence, in order to improve the penetration of ultrasound in the tissues of human body a coupling gel is used.

Рис. 3. Reverse scattering of ultrasound.

At oblique incidence of the ultrasound beam, an angle of incidence, angle of reflection and angle of refraction are determined (Fig. 2). The angle of incidence is always equal to the angle of reflection. The ultrasound beam is necessarily refracted when it intersects the interface between two media with different speeds of ultrasound conduction. The ratio of the sine of the angle of refraction to the sine of the angle of incidence is equal to the quotient of the speed of propagation of ultrasound in the second environment to that in first. The greater the difference between the speeds of ultrasound in both media the stronger is the refraction. There is no refraction if the speed of ultrasound is the same in both media or the angle of incidence is equal to zero. The described so far reflection of ultrasound is called mirror reflection. It takes place when the wavelength is many times greater than the size of the inhomogeneities on the reflecting surface. In case the wavelength is comparable with the inhomogeneities on the reflecting surface or inside the reflecting medium the ultrasound is scattered. When the ultrasound is scattered in the same direction whence it comes we speak of backward scattering (Fig. 3). The intensity of scattered ultrasound increases with the number of inhomogeneities of medium and with the frequency of ultrasound. The scattering depends slightly on the direction of the incident beam that allows better visualization of reflective surfaces and media, such as the parenchyma of organs.

Fig. 4. Electronic multielement probes - linear (left), sectoral (middle) and convex (right) emitters.

When passing through a homogeneous medium, the ultrasound decreases its amplitude because a part of its acoustic energy is absorbed and converted to heat. If the medium is heterogeneous, the ultrasound amplitude will be additionally reduced due to the reflection and scattering. The total weakening, or attenuation, of ultrasound is characterized by the so called damping ratio, this is the weakening of the ultrasonic signal per unit distance. Attenuation is measured in decibels (dB), and the damping ratio - in decibel per centimeter (dB/cm). The damping ratio increases with increasing frequency.

Emitters and sensors of ultrasonic waves

In medicine, ultrasound is generated by the application of short electrical pulses to a piezoelectric crystal. Under the influence of alternating electric voltage the piezocrystal (usually lead zirconate or titanate) deforms (inverse piezoelectric effect) producing a mechanical wave. Detection of reflected signals is based on the direct piezoelectric effect (the deformed crystal produces electric voltage). Usually the ultrasonic device uses the same piezoelectric probe (transducer) for both emitting and detection of ultrasound waves.

Fig. 5. Multielement phased array of piezocrystals for electronic focusing and steering of emitted ultrasound waves.

Modern ultrasound transducers contain a large number of piezoelements arranged in a linear way (Fig. 4). Depending on the arrangement we distinguish linear, sector and convex type of transducers. To build a two-dimensional image on the screen of sonographic instrument the ultrasonic beam is repeatedly directed from the left edge to the right edge of scanned plain of human body. Beam steering in different directions, ie, the sweeping of beam direction laterally, is achieved electronically submitting the electric excitation to the piezoelectric crystals of multielement emitter in different phases (Fig. 5). Simultaneously and in the same manner the beam is focused at the required depth in the tissue. The emitted ultrasound represents a complex ultrasonic field, comprising a near area, a far region and intermediate focus region (Fig. 6). The place of greatest narrowing of the ultrasound beam is called focal area (focus) and the distance between the emitter and the focus is the focal length. The quality of the image is highest just in the focal zone. 2D-sonography uses pulsed ultrasound, i.e. few cycles of oscillations separated in time with gaps of no vibration. Each bunch of oscillations is called ultrasonic pulse (Fig. 7).

Fig. 6. Zones of ultrasonic field focused at a certain distance from the front surface of the ultrasound emitter.

Pulse duration is the time, filled with a single pulse. It usually contains 2-3 cycles of oscillations. Pulse-repetition period (PRP) is the time in μs between the onsets of two neighbour pulses. Pulse repetition frequency (PRF) is the number of pulses emitted per unit time (1 s) and is usually expressed in kilohertz (kHz). To build an image 2D-sonography uses very short ultrasound pulses emitted usually 1.000 times per second (pulse repetition frequency 1 kHz), whereat 99.9% of the time the piezocrystal works as a receiver of reflected signals. The duty factor is another important concept for pulsed ultrasound - this is the portion of the total time in which the ultrasound is emitted in the form of pulses. Spatial pulse length is the space length occupied by a single ultrasound pulse (Fig. 7). Spatial pulse length and the pulse duration both decrease if the number of cycles in the pulse is decreased or if the frequency is increased.

Fig. 7. Spatial pulse length.

Example: Pulse duration is 4 µs, PRP is 160 µs. Hence, the duty factor (dimentionless) = 4 / 160 = 0.025 = 2.5 %. Typical duty factors for sonography are ~ 0.1 to 1.0 % and for Doppler sonography are ~ 0.5 to 5.0 %. The spatial resolution of sonographic imaging is defined as the ability to distinguish two points as seperate in space. Spatial resolution is usually measured in mm. The higher the spatial resolution, the smaller the distance which can be distinguished. Spatial resolution is further categorized as axial resolution and lateral resolution. Lateral resolution is defined by the minimal distance between two points, situated perpendicularly to the ultrasound beam, still separated on the screen. It is equal to the diameter of the ultrasound beam, so the well resolved image could be formed using a narrow, focused beam. If the beam is diverging, the lateral resolution becomes poor with increasing the depth of penetration. Axial (longitudinal, azimuthal) resolution is represented by the minimum distance between two points along the axis of the beam, which are still resolved on the screen. It equals the half of the spatial pulse length, hence the shorter the pulse the better the axial resolution. As a rule, the axial resolution is better than the lateral one.

Main scheme of 2D-sonographic instruments

Fig. 9 displays the basic blocks of a modern ultrasound instrument. Electronic generator sends electrical impulses to the ultrasound probe, which produces short ultrasonic pulses in the direction of the examined tissue, receives the reflected echos and converts them into electric current pulses. This current is too weak and is fed into a radio frequency amplifier. The gain of this amplifier is controlled by a separate block called Time-Gain Compensator (TGC) in order to account for tissue attenuation. When ultrasonic signal penetrates into tissue it attenuates exponentially with the depth, hence the echoes will be as weak as later they come back to the receiver. Hence, the brightness of the images on screen should strongly decrease with increasing the depth at which the respective objects are located. The Time-Gain Compensation block measures the delay time of the reflected ultrasonic signal and changes the gain of the amplifier so that a longer delay corresponds to a greater amplification. Thus, this block allows obtain an image on the screen with the same brightness independently on the depth. The amplified electrical signal is further rectified and filtered by a demodulator block. Finally, the current is amplified by the videoamplifier and is fed to the monitor.

Fig. 8. The main circuit of a 2D- sonoechograph

The entire image of the observed plain section is called a frame. The frame is formed by a large number of vertical lines (Fig. 9), each one representing the path travelled by a separate ultrasound pulse. Thus, the time to scan one frame is equal to the pulse duration x the number of scan lines per frame. The number of frames reproduced in 1 s is called frame rate (Hz). Typical frame rates in echo imaging systems are 30-100 Hz. When the observed object is moving, for example the heart, a temporal resolution is defined to assess the ability of medical ultrasound instrument to detect the movement of the object over time. The temporal resolution is equal to the frame rate. Thus, the temporal resolution or frame rate = 1/(time to scan 1 frame).

Fig. 9. An image of a frame on the screen. Each line represents the path traveled by a single ultrasonic pulse.

Pulse repetition frequency (PRF) confines these parameters and predetermines the quality of sonographic imaging. This is due to the obvious relationship between the pulse repetition frequency (PRF), the number (N) of lines forming one frame, and the number of frames per second (frame rate): PRF = N × frame rate. In modern black and white sonographic imaging apparatus the pulse repetition frequency is 1 kHz (1000 pulses per second). In modern instruments the quality of image is increased using composite frames instead of simple frames. When the observed object is placed far from the emitter a single frame is sufficient, provided the ultrasound pulses are focused on this object. In medical imaging the system must capture both close-in reflections in the nearfield, such as in vascular scans, as well as reflections in the farfield for abdominal or cardiac clinical applications. To achieve focusing at different scan depths, the ultrasonic instrument uses multiple sweeping of the observed cross-section. In the first sweep the beam is focused on the closest layer of tissue, and the image obtained from this zone is saved. On the next sweep the focus is moved in the next zone obtaining the second image which is again stored in the device memory, etc. The result is a composite image, focused on the whole depth, called a composite frame. This way of working and this type of instruments are called dynamic focusing, respectively, ultrasound instruments with fast scanning in real time. Since the build up of a composite frame requires considerable time, the number of received frames in 1 second (frame rate) is lower. The frame rate is reduced by a factor equal to the number of focal points used per scan line. If there are moving tissues in the observed cross section the number of frames per second must be high in order to represent truly its movement on the screen. A frame rate of at least 20 frames per second is needed to give a realistic illusion of motion.

Fig. 10. A-mode and B-mode representations of the echo signals along a single scan line.

The image of the screen of 2D-ultrasound instruments has a digital (discrete) type and is formed by a large number of points (pixels), distributed in a matrix of 512 × 512 pixels. Typically, a pixel corresponds to the observed area of about 0.4 mm2. The brightness of each pixel is represented as a number of degrees of gray, usually 16-32-64- 128-256 degrees (4-5-6-7-8 bits).

Fig. 11. М-mode representation of the echo signals, reflected from moving objects, along a single scan line.

The earliest ultrasound imaging instruments used one-dimensional representation of the reflected ultrasound signal. To scan the depth of object they used a single beam directed along a fixed axis. The beam was displayed as a single scan line on the screen with the reflected echoes distributed on the line depending on their time delay. The intensity of each echo signal was represented as peak height (A-mode) or as the brightness of a point (B-mode), Fig. 10. To detect the movement of a reflecting structure, M-mode was used in which each point of the scan line is represented as perpendicular line depicted over time (Fig. 11). If the reflecting structure (heart valve) is moving the perpendicular line is wave- like, otherwise it represents a straight, solid line.

Fig. 12. Examples of sonographic images: two- dimensional type B (left), Type M and cardiogram (right).

The 2D-mode is the default mode that comes on when any ultrasound / echo machine is turned on. It is a twodimensional cross sectional view of the underlying tissues and is made up of numerous B-mode scan lines. A portion of the tissues that are intersected by the scanning plane is displayed on the screen. Depending on the probe used, the observed field could have a sector shape (for abdominal organs – Fig. 12) or rectangular or trapezoid shape (for superficial or vascular observations). On a grey scale, the high reflectivity tissues (bone) are white; low reflectivity tissues (muscles) are grey and places with no reflection (water) are black. Deeper structures are displayed on the lower part of the screen and superficial structures on the upper part.

2. PHYSICAL BASIS OF DOPPLEROGRAPHY IMAGING

Doppler effect is the basis of all doplerographic methods for ultrasound diagnostics (doplerography). The effect represents a shift in the frequency of the received acoustic wave from that of the emitted one when the emitter and receiver are moving relative to one another, or, when the received wave is reflected from a moving object. The latter case is typical for the medical applications of the effect where, the reflecting bodies - red blood cells, heart walls and valves are constantly moving. Relatedly, the Dopplerography is used for the study of the moving tissues and media in human body. Depending on the mode of the image acquisition and data processing the doplerographic diagnostic methods can be divided into: 1. Continuous wave Doplerography – used for precise measurement of the spectrum of velocities of blood flowing in a vessel and in the heart. 2. Pulsed wave Doplerography (spectral Doppler ultrasound; D-mode) – used to locate the blood flow and measure the spectrum of velocities of blood flow. 3. Color Doppler ultrasonography. This includes two-dimensional color imaging of blood flow, in which the speed of the individual elements are marked with colors and their shades (color flow mapping, power Doppler) and others.

Doppler shift in frequency. Doppler angle

The diagnostic ultrasound devices, based on the Doppler effect, do not measure the intensity nor the frequency of the received signals; they just determine the difference between the frequency, fo, of the emitted signal and the frequency, f, of the received signal, ie, the frequency difference fD = fo - f, called Doppler frequency shift. This is carried out by the doplerographic instrument which superimposes the emitted signal and the received one (Fig. 1). Based on their phase difference it calculates the Doppler frequency shift, fD, and hence, the velocity of reflecting object.

Fig. 1. Phase difference, Р, between the transmitted (black) and the received (red) signal.

The sound speed, C, in soft biological tissues averaged 1540 m/s. The speed, V, of the moving biological structures (e.g., blood) rarely exceeds 1.5 m/s (in the highly constricted vessels it reaches to 7 m/s), i.e., it is much smaller than C (V << C). Let us denote by α the angle between the velocity vector of the reflecting surface (direction of movement) and the direction to the probe (Fig. 2A). This angle, α, is typically called a Doppler angle. The Doppler frequency, fD, is determined by the projection of the velocity vector, V, on the line linking the reflecting surface and the probe, i.e., by the product V.cosα in the main equation: fD = 2fo (V/C) cos α. Using this formula, the velocity, V, of the moving tissue could be calculated provided the Doppler angle, α, and the Doppler frequency, fD, are determined and knowing the frequency, fo, and the speed, C, of emitted sound (Fig. 2B).

When the reflecting surface moves towards the probe, the Doppler frequency shift is positive (fD > 0). If the reflecting surface is moving away in the opposite side of the probe, then the projection of velocity, Vcosα, is negative and the Doppler shift is negative (fD <0).

Fig. 2 A. Doppler angle, α, between the direction in which the reflecting object moves and the axis to the probe.

The above indicated equation shows that the Doppler frequency shift, fD = 2fo (v/C) cos α, is proportional to the frequency, fo, of emitted signal. For this reason, it is appropriate to use higher frequencies of the signal. Unfortunately, with increasing the frequency the absorption and attenuation of ultrasonic vibrations also increase reducing the depth of the ultrasound penetration in medium. Therefore, compared to the B-mode, the pure Doppler mode uses slightly lower frequency signals as follows: • 2 MHz - examination of the brain vessels; • 3 MHz - examination of the blood flow of the placenta; • 4 or 5 MHz - examination of relatively large and deeply seated vessels; • 8 or 10 MHz - examination of small, superficial peripheral vessels.

Fig. 2 B. Calculation of the speed of the blood based on the Doppler frequency shift.

Doppler frequency shift has values that lie in the audible frequency range of 16 Hz - 20 kHz. All sonographic instruments with Doppler mode employ proper electronic amplifiers and loudspeakers to transform the Doppler frequency shift into audible sounds which the diagnostic doctor uses in his practice.

Positive and negative Doppler frequency shift

Very often the doctor needs to know only the projection of the speed of blood, V.cosα, and its change over time. The Doppler angle, α, can be evaluated based on the dependence of the Doppler frequency shift, fD, on the angle, α (Fig. 3).

Fig. 3. Dependence of the Doppler

frequency shift, fD, on the Doppler angle, α, between the axis of the ultrasonic probe and the direction of blood flow. Doppler shift, fD = 0 if the axis of the probe is perpendicular to the blood vessel, respectively, the angle α = 90° and cosα = 0. In this case it is not possible to measure the velocity of blood flow. However, it is sufficient to tilt the probe to another known angle, measured in respect to the line, as established to be perpendicular to the vessel, and it becomes possible to assess the flow velocity based on the value of the fD. When α <90° the blood is approaching the probe and fD> 0. When α> 90° the blood flows away from the probe and fD <0. To increase the accuracy it is useful to orient the probe so that the Doppler angle, α, will have smaller values. However, if the probe is excessively inclined with respect to the blood flow and the angle, α, becomes smaller than 25° or greater than 155° (the so-called critical angles) the ultrasound might be completely reflected on the vessel wall and the measurement will become impossible.

Spectrum of velocities of the blood flow

The average speed of blood flow in the arteries is not constant, as it varies periodically in accordance with the cycles of contraction-relaxation of heart. Therefore, the Doppler frequency shift, fD, will also fluctuate over time. For example, if the heart rate is 150 beats per minute, the period (cycle) of the heart beatings will be Tc = 0.4 s. In order to detect any change in the blood flow velocity at different points of the cardiac cycle (sistolic and diastolic) we need at least 10 equally displaced measurements within this time interval. This means that the measurements should be conducted at intervals of not more -1 than Tc/10 or 0.04 s, i.e with a repetition frequency of not less than 25 s . In this case the requirement for real-time measurement is fulfilled.

Fig. 4. Parabolic distribution of blood flow velocity across the cross section of blood vessel.

The normal blood flow through the blood vessels is laminar, and the speed of separate layers decreases from the center to the walls of the vessel in a parabolic law (Fig. 4). Therefore, for a fixed moment of time, the reflected echo signal will contain components with different Doppler shifts. Each component will be reflected from a particular layer moving with its own speed. The distribution (spectrum) of speeds of different layers across a typical blood vessel as dependent on cardiac cycles is shown in Fig. 5. Greater intensity will have the echo signal reflected by a larger number of layers, moving at the same speed. These signals are shown as vertical lines with higher brightness. On the other hand, the greater the speed of the layer, the greater will be the Doppler shift of the signal reflected from it. During the systole the spectrum is shifted to the right, and during the diastole – to the left of the velocity axis.

Fig. 5.

Distribution of velocities, V, in various layers of a blood vessel and of their respective Doppler shift, fD, at the phases of systole (A) and diastole (B).

The shape of the spectrum of speeds highly depends on the geometry of the blood vessel. In wider vessels most of the layers have a same speed hence, the spectral peak will have a small width (Fig. 6A). The narrow vessels exhibit greater resistance and friction, their layers move at speeds that are more evenly distributed and their spectrum will be wider (Fig. 6 B). At stenosis there are vortices within the vessel in which the speed is opposite to the speed of mainstream. Therefore, the spectrum will be even wider and echo signals with negative Doppler shift could appear (Fig. 6 C). Similar broadening of the spectrum of speeds also occurs at each curvilinear section and bifurcation, where the blood vessel bends or branches out at different directions. In both cases, the velocity of the blood changes its direction and there appear layers moving at different angles relative to the direction of the incident beam. The width of the incident ultrasound beam also affects the width of the spectral peak increasing it.

Fig. 6.

Velocity distribution in vessels with various diameters. А – constriction of the specter for a broad vessel, B – widening of the specter for a narrow vessel, and C – show up of turbulence and negative peak for a vessel with stenosis.

The sign of the Doppler signal, respectively, the direction of the spectral peak will depend on the direction of blood flow towards or away the probe (Fig. 7). The systole - diastole alternations make the walls of large blood vessels (without veins) to oscillate with low frequency. These oscillations are due to the elasticity of the walls and contribute with an additional component in the low-frequency part of the velocity spectrum of blood. This component is usually removed by means of low-pass filters that cut off all frequencies from zero to about 80 or 120 Hz. When observing the blood flow in veins this filter can be excluded. Further distortion of the spectrum occurs due to the following two reasons. 1) Due to their small size, the erythrocytes better reflect signals with higher frequency, which results in a shift of the spectrum to the high frequencies. 2) On its passage through the tissues, the ultrasound is stronger absorbed and attenuated if its frequency is higher. This leads to a shift of the observed spectrum to lower frequencies, especially at deeper penetrations of ultrasound.

Fig. 7. On the left: velocity distribution as dependent on the direction of motion of the blood relative to the probe. In the center: a spectrum of velocities in laminar and turbulent flow of blood. Right: echogram of a blood vessel. The horizontal axis indicates the time and the vertical axis - the Doppler frequency shift. Basic kinetic parameters of blood flow

Based on the shape of the Doppler (frequency) spectrum and using a particular algorithm we can calculate some important kinetic parameters of blood flow (Fig. 8). These are the maximum (systolic) speed, A, the minimum (diastolic) speed, B, and the average speed, M. These values, however, depend on the Doppler angle. They could be corrected for the large straight line blood vessels, whose Doppler angle is easily determined by anatomic considerations. However, such a correction is a problem for smaller, curved vessels (e.g. vessels in a tumor mass) as the assessment of their Doppler angle is difficult. To solve this problem several calculations, independent on the Doppler angle, are proposed as listed below. Systolic-diastolic ratio (peak systolic to end diastolic ratio: SDR = A / B. Resistance Index (resistance): RI = (A-B) / A. Pulsatility Index: PI = (A-B) / M. Percentage Stenosis (PS). It is used to assess the degree of vessel (artery) contraction. For its calculation the maximum blood speed (Fig. 9) is measured in an area prior to stenosis, VPS, and within the zone of stenosis, VS, using the formule: PS (in %) = 100 (VS - VPS) / VS.

Fig. 8. Doppler spectrum of blood. A - systolic velocity, B - diastolic velocity, M - average speed of blood.

Using the Doppler measurements, we can also calculate other important quantitative characteristics of the blood flow: stroke volume of blood, the area of mitral orifice, blood volume flux, kidney aortal index, pressure gradient, time for acceleration of the flow, time delay of the blood flow, linear velocity integral.

Fig. 9. Speed of blood prior to the area of stenosis, VPS, and within the area stenosis, Vs.

Continuous wave (CW) Doppler

To obtain a Doppler picture of the tissues the following two types of techniques are mainly used - continuous wave Doppler (CW-Doppler) and pulse wave Doppler (PW-Doppler) echography. CW-Doppler is the older and electronically more simple of the two kinds. In CW-Doppler the probe comprises two crystals, one of which constantly emits ultrasound and the other receives the reflected signals and sends them for further processing. The receiving crystal measures the Doppler shift of the reflected beam, which is usually a thousandth part of the frequency of the emitted ultrasound. After amplification the signal, corresponding to the Doppler shift, is transmitted to two speakers and, at the same time, to the monitor to make qualitative and quantitative assessment of the observed moving tissue. In heart echography, the duration of the emitted and reflected echo signal is not infinite and is limited to durations not exceeding 5-10 ms. This is related with the need to evaluate the change in the average speed of blood flow during the various phases of cardiac cycle, i.e. to realize the principle of "real-time measurement". The block diagram of continuous wave Doppler echograph (Fig. 10) includes electronic generator whose pulses are fed to the emitting crystal which converts them into continuous ultrasonic beam with a frequency, fo. The echo signal, reflected from the examined tissue, with frequency, f, is detected by a separate crystal and converted to electrical signals which enter the main processing block. Most favorable conditions for tissue examination are met in the area (volume) where the emitted and received ultrasound waves overlap. This is so called controlled area (controlled volume). The vast majority of ultrasound echoes are reflected by stationary discontinuities located in the controlled volume. They have the same frequency, fo, as that of the emitted ultrasound and these signals are removed and not processed. If the controlled volume contains a moving structure (blood, heart walls and valves) then the reflected echo will contain a signal with frequency f, different to fo. What is processed is namely this component, which is amplified, and the magnitude and sing of its Doppler shift, f D, determined. The processing unit further generates an alternating current with the same frequency of fD and sends it to one of the two loudspeakers depending on the sing of the Doppler shift. One of the speakers receives signals corresponding to the blood flow which approaches the emitter (+ sign of fD) and the other one receives the signals for receding blood flow (- sign of fD). Simultaneously, this block sends the magnified output current to the spectral analyzer, which derives the velocity spectrum of reflected echo signal. Signals with Doppler frequency shift can be stored and viewed on screen as a graphical distribution on frequency (Fig. 7).

Fig. 10. Block diagram of Doppler echograph with constant wave.

Specific for this type of Doppler echography is the large controlled volume of examination. This makes it difficult to analyze the total reflected echo signals especially if this volume contains more than one blood vessel. The large volume under control, however, is an advantage when testing the valvular stenosis in heart and calculating the pressure drop on heart valves. The continuous wave Doppler echography is applied mainly to study the blood flow in peripheral vessels and for the analysis of the atrioventricular and aortal blood flow.

Fig. 11. A series of emitted packages of ultrasonic vibrations (a) and the delay, Δt, of their reflected echoes (b).

The main advantage of CW Doppler is its ability to measure high blood velocities accurately. The main disadvantages of the method are: • The spectrum of speeds could be obtained only if the controlled volume contains a single blood vessel. All speeds of the blood layers within the blood vessel are precisely determined, however, the location of the blood vessel could not be determined. • The information obtained corresponds to the entire depth of the controlled volume. In many cases the controlled volume contains several blood vessels with different directions and speeds of blood. Due to the absence of axial space resolution it is impossible to examine each blood vessel separately.

Pulsed wave (PW) Doppler echography (spectral Doppler)

The above disadvantages are removed in the Doppler echography which uses ultrasound oscillations grouped into wave packages, well separated from one another. This type of ultrasonic instruments has almost the same structure as that of the continuous wave scanners (Fig. 10) to which a special new block is added. The function of latter block is to periodically quench the electrical impulses, generated by the electronic generator. The time intervals when the electric oscillations are quenched determine the so called repetition period, T (Fig. 11). The oscillations generated outside the quenching period fed the piezocrystal that emits packages of ultrasonic vibrations and receives their reflected echoes. When the package of oscillations is reflected by a certain layer it returns to the piezocrystal with the retention, Δt, which depends on the distance (depth) between the emitter and the reflecting layer (Fig. 11). To allow the piezocrystal to detect the reflected packages of oscillations it is obligatory that Δt << T. In fact, the piezocrystal receives as many echoes as is the number of reflecting layers in examined tissue. Each of these reflected packages has its one delay corresponding to the depth of its reflecting layer. In order to allow the scanner to monitor only one of these reflective layers a fixed time delay (called a time door or window) is assigned in the main processing block, which makes the scanner to process only the packets reflected from a certain depth (control volume). Thus, the observed tissue is divided into narrow sections (controlled volumes or strobes), and each of them can be tested by setting the appropriate level of time window.

Fig. 12. Unlike continuous wave doplerography, the pulsed wave doplerography can test each layer separately.

There is an upper limit to the Doppler frequency shift, respectively, to the determined velocity (≤ 2.0 m/s), which can be detected by the ultrasonic pulsed wave echographic instruments. This limit is called Nyquist limit and is equal to half the frequency of pulse repetition (1 / T). When this limit is overrun the Doppler spectrum becomes distorted and the determination of velocity is highly incorrect. This condition is called aliasing. Usually, the thickness of the controlled volume is about 5-10 mm, and it can vary as desired. It is equal to the spatial length of the package of oscillations. So the attention can be drawn on a separate blood vessel, heart valve or the wall of heart (Fig. 12). When the volume control is sufficiently small, the blood velocity distribution can be investigated in a separate vessel.

Ultrasound echography with color Doppler mapping

Modern apparatus for spectral Doppler possess a regime for two-dimensional color imaging of moving tissue, referred to as color Doppler mapping. As in the regime «spectral Doppler", in this mode the probe emits ultrasonic packages of oscillations (pulses) and receives the reflected signals coming from different depths. The depth of each reflecting structure is determined by the delay time of the return pulse, because the speed of ultrasound is almost the same throughout the soft tissues. Signals reflected from a stationary object have the same frequency as the frequency of emitted signals while signals reflected from moving objects contain a Doppler frequency shift. As a rule, the signals reflected from stationary objects are much stronger, and those reflected by moving objects (such as blood in a blood vessel) have much less intensity (sound energy). In the conventional Doppler echography the reflected signals that do not contain a Doppler frequency shift are suppressed and do not appear on the screen. Conversely, in ultrasonography with color Doppler mapping these signals are displayed on the screen, but in gray or in black and white (Fig. 13). The brightness of a point on the screen, corresponding to such a reflected signal, is proportional to the amplitude of the echo signal as in the B-mode echography. If any echo signal, corresponding to a given element on the screen, contains Doppler frequency shift, the apparatus determines the magnitude and sign of this shift and displays them in color (Fig. 13). The red color indicates the approaching of the reflecting object to the emitter, and different hues of color (from dark red through red to orange and yellow) show the average speed value. If the movement is in opposite direction a blue color is used and again its hues (from deep violet and almost black to blue) show the average speed. In addition to the direction and average speed of blood flow, some types of instruments display the width of the speed spectrum in the blood vessel. One way to display it is to change the color saturation of the base color. For this purpose, the base color is mixed with white color in different proportions, the broader the velocity spectrum, the more white color is added, respectively, the color saturation is reduced.

Fig. 13. Formation of a two-dimensional image with color Doppler mapping. On the left - echo signals reflected from various reflecting objects: stationary and moving to or from the emitter of ultrasound. On the right – color imaging of the observed structures on screen.

There are several ways of displaying the movement of examined tissues: mapping of the relative speed and direction of movement of tissues; mapping of the acceleration (change of velocity over time) of the moving tissues and others.

Fig. 14. Color image of heart valve obtained by Power Dopple (PD). Fig. 15. Display of the movement of heart wall by the method of tissue Doppler imaging. Two-dimensional (above) and one-dimensional (below) modes. The color Doppler mapping (CFM) is surerior to the spectral Doppler mode (CW or PW) as it is able to examine the blood flow in a large controlled volume and in many blood vessels in real-time, i.e., with frame rate not less than 15 ÷ 20 c-1. This is achieved increasing the spatial and time resolution by increasing the duration and number of oscillations in each package of oscillations to between 8 and 24. The frame rate is also increased to several tens in 1 second. These images are stored and then viewed in slow speed. This way of storing and viewing images is called cine memory, or cine loop. Another variant of color Doppler mapping is the so called Power Doppler (PD). What is determined in this mode is the energy, not the frequency, of the Doppler frequency shift of the reflected signal. This approach allows to increase the frame rate and, thereby to increase the sensitivity to lower speeds and to reduce the dependence on the Doppler angle. These advantages, however, are accompanied by the loss of ability to determine the value and direction of the blood flow velocity. The method only determines the availability of motion. It is conveniently used to monitor the movement of heart valves (Fig. 14). Another mode is the three-dimensional color Doppler mapping. It displays a three-dimensional picture of how the blood flow velocities are distributed in space and gives its evolution over time. The difficulty with this method is to obtain not less than 20 ÷ 30 frames per second to produce an allusion of movement in real time. For example, a cardiac cycle of contraction-relaxation lasts about 1 second. During this time, the heart is scanned in a two-dimensional plane which is moved in the third coordinate of space. If the number of received frames per second is small the movements of the walls will be displayed in a distorted form. These technical difficulties can be overcomed by reducing the controlled volume or by simultaneous reception of reflected signals emitted by several piezocrystals.

3. APPLICATION OF LASERS IN MEDICINE

Laser light is conveyed to the targeted tissue in following ways: 1) Through optical cables that contain flexible quartz fibers. 2) Through hinged tubes that contain mirrors at its ends. 3) By manually directing the beam to the tissue. Much greater accuracy is achieved when the targeting is carried out using a micromanipulator under a microscope. Further improvement of the accuracy may be achieved using a computer for automatic scanning of the object and directing the laser beam.

1. Effects of laser light on tissues

The interaction of the laser light with cells and tissues is determined by the optical properties of the cells and tissues and the wavelength of light. Laser light, like other types of electromagnetic radiation exerts photothermal, photomechanical and photochemical effects on tissues. Photothermal effect is due to the heat generated by the laser light in irradiated tissue. If the tissue is heated to a temperature below 43 °C, the damage caused is reversible. The damaged tissue can recover even after a very short heating to 50 °C. Higher temperatures (50 - 100 °C) cause denaturation and coagulation of proteins, which leads to irreversible thermal damage of tissues. At temperatures close to 100 °C the tissue water evaporates quickly. Temperatures around 170 °C cause carbonization of the tissue. Such intense warming is caused by lasers with continuous or pulsed operation, if the laser beam has very high intensity, or energy flux density. Specific targets (particular parts of tissue) undergo selective photothermolysis when they absorb a laser beam generating such an amount of heat that overwelmes the natural cooling of the tissue. The basic condition for obtaining selective photothermolysis is to use laser light of such wavelength that is strongly absorbed by certain molecules of the tissue, called chromophores. Each chromophore absorbs the laser light and converts it into heat, which continuously accumulates due to the arrival of new portions of light before the energy of older ones dissipates. This is achieved when the laser beam consists of powerful pulses whose repetition period is less than the time for heat relaxation of the target. The latter is a measure for the natural cooling rate of the target and is equal to the time required to lower the temperature of the target to half of that achieved immediately after exposure. This time is generally between 300 and 700 s and it is shorter in case the target has smaller dimensions. Selective photothermolysis results in the increase of local temperature selectively impairing only those tissues which contain chromophores with minimal damage to surrounding healthy tissue. To achieve this effect three basic requirements must be met: 1. Proper selection of a suitable wavelength, whereat the targeted chromophores exhibit specific absorption. 2. Proper selection of a suitable repetition period of the laser pulses shorter than the cooling time of chromophores. 3. Proper selection of a suitable intensity of the luminous flux in order to reach the critical high temperature of the target.

Fig. 1. Absorbance spectra of water, hemoglobin, and human epidermis.

The main chromophores of cells, absorbing ultraviolet light in the C and B bands (< 320 nm) are melanin, hemoglobin, collagen and other proteins, as well as bilirubin, urocanic acid and nucleic acids. Melanin absorbs strongly in the region of 320 - 1000 nm. Water is the predominant absorber of radiation with wavelengths exceeding 1000 nm (Fig. 1). Additionally, subcutaneous collagen bundles scatter strongly the light with wavelength more than 300 nm. UV light from the area A, as well as the blue, green and yellow lights is strongly absorbed by the blood hemoglobin and plasma bilirubin. Optical region between 600 - 1300 nm is referred to as "optical window" (Fig. 1) because the light with such wavelength easily penetrates deep under the skin without significant absorption and scattering. Melanin is the main chromophore in this area. In case the laser light has very high intensity, its photothermal effect turns into a photomechanical one. Under such intense very thin laser beam the irradiated targets are dispersed due to their thermal expansion and evaporation. In addition such effect takes place during the application of super strong selective photothermolysis. Photochemical action of the laser takes place when tissues are irradiated by low intensity light of unfocused laser and the light is resonantly absorbed by specific chemical groups, chromophores, associated with tissue biomacromolecules or introduced from outside. This can lead to various photobiological effects. Special case of photochemical action represents the so called photodynamic therapy, wherein the chromophore groups generate free radicals damaging the irradiated cell.

2. Examples of application of laser technology in medicine

Based on the main effects of laser light the medical application of laser technology is divided into three major groups: 1. Application of thermal effect in surgery – laser scalpel (excision, hemostasis and evaporation of tissues); 2. Selective photothermolysis of superficial vascular disorders, pigmentation disorders and tattoos; 3. Photodynamic therapy of tumors. Thermal effect is produced in tissues through irradiation by a focused laser whose intensity has medium to high levels. This effect is highly pronounced in pigmented tissues, strongly absorbing laser light, such as tumor tissues that contain more melanin. Depending on the depth of penetration and the thermal conductivity of the irradiated tissue, the laser beam can vaporize the surface layer, or can cause local destruction of the underlying layers due to the thermal denaturation and coagulation of proteins. On the other hand, coagulation of proteins by laser beam causes the gluing of damaged blood vessels - haemostatic effect. Based on these effects the thermal effect is used in ophthalmology to reshape the cornea eliminating the corneal astigmatism and strabismus, to drill hole reducing the intraocular pressure in glaucoma. In surgery, the thermal effect of laser beams is used to open blocked blood vessels, to break up kidney stones (laser lithotripsy) and to remove tumor tissue by precise laser scalpel, working mainly with infrared light. In many cases, the laser beam is directed through a flexible fiber that is inserted into a blood vessel or body cavity through a cannula. In laser oncosurgery and microsurgery, the beam of laser scalpel is directed to hardly accessible areas using flexible optical fibers. This is called laser ablation, i.e., sterile removal of aberrant tissue by laser beam. It is greatly affected by the ability of tissue to absorb energy and by the intensity of laser beam.

Fig. 2. Penetration depth of laser beams into skin, as depending on the wavelength (nm).

Specific tissues can be selectively damaged by the heat produced by laser light whose wavelength corresponds to the absorption of the characteristic chromophores of tissue (Fig. 1 and 2). The light of ruby laser (694 nm) or Nd:YAG laser (1064 nm) penetrates deep into the skin and is absorbed by melanin containing cells, leading to their selective photothermal lysis. The light of argon laser, as well as the yellow light of some lasers is strongly absorbed by hemoglobin which is used for the treatment of vascular damages. The light of CO2 laser has a wavelength of 10600 nm and is mainly absorbed by tissue water hence, it poorly penetrates into the skin and does not have a selective action. Selective photothermolysis is used to stimulate the skin fibroblasts to synthesize more collagen which has a rejuvenating and healing effect on the skin. The enhanced collagen synthesis continues for about 2 months and maintains a high level of collagen for about 6-12 months. This results in the removal of skin wrinkles because the thickness of collagen layer increases. The laser light is conveyed to the skin through a sapphire crystal which closely abuts the skin. The role of the sapphire is to take up the excess heat from the skin, since sapphire has a 30 times higher thermal conductivity compared to water. This is an example of photothermal stimulation, in which the fibroblasts are heated briefly to 60°C, which stimulates them to produce collagen. It is believed that this process is similar to the mechanism of wound healing by selective stimulation of fibroblasts.

Fig. 3. Outside view of lasers used in medicine: a) in the form of pencil; b) lasers with adjustable power; c) lasers with digital control.

Vascular injuries are treated by a powerful laser beam with a wavelength of 535 - 615 nm which is selectively absorpbed by hemoglobin. When the temperature reaches 60 – 65 °C the blood coagulates causing blockage of the vessels of medium size. In this treatment the light having wider range of wave lengths is preferred, as the monochromatic light is absorbed only to a certain depth causing an inhomogeneous heating of the tissue.

Fig. 4. Outside view of lasers used in medicine: a) an optical head in the form of a bell; b) laser with a permanent magnet; c) lasers with a pencil; d) lasers with infrared and ultraviolet light

Selective photothermolysis is used to fix a detached retina. Suitable pulsed beam of laser light is transmitted through the transparent layers of the eye (cornea, crystalline lens and vitreous body) without damaging them. Reaching the highly pigmented retina the beam is absorbed there causing heat denaturation of local proteins - photocoagulation. Laser beam exerts selective effect on different organelles within the cell; those who absorb it at the selected wavelength are destroyed while the transparent parts of the cell remain unaffected. Chloroplasts and mitochondria can be selectively tested in this manner. Thus, by adjusting the laser at different wavelengths one can vary the degree of absorption of laser light by certain cells or cell organelles. Selective photothermolysis is used to remove tattoo. This is done by high-energy pulses of ruby laser, which reach the ink particles of the tattooed skin and disperse them. Photochemical effect is used to treat skin diseases, to accelerate healing of wounds and bone fractures and for treatment of inflamed mucous membranes and gums in dentistry. Tumor cells specifically absorb some endogenous, photosensitive substances (photosensitizers) such as porphyrins from the blood. This is used in photodynamic therapy (PDT) of tumors, which is based on the photochemical effect of laser light on these photosensitizers. This treatment uses low-power lasers that emit in the visible region: argon laser emitting at 480 - 515 nm, helium-neon laser emitting at 632.8 nm, and krypton laser emitting at 521 and 530 nm (table 1). Light emitted from these lasers is absorbed by the photosensitizers (endogenious tissue pigments - hemoglobin, melanin) and excites their molecules. Upon irradiation of tissue by argon laser, the accumulated porphyrin passes into excited singlet level and fluoresces red light, which indicates the presence of a tumor cells. Irradiating the tissue with helium-neon laser excites the porphyrin molecules into triplet level which represents highly reactive chemical radical, killing the host cells, i.e., the tumor cells. Very often synthetic photosensitising substances, also accumulated selectively in tumor tissue, are used as a treatment modality. Medicine uses different types of lasers (Table. 1) which, depending on the physical state of their active medium may have special characteristics and advantages. The solid-state lasers include ruby, neodymium-yttrium (Ne: YAG) laser and alexandrite laser. Gas lasers include carbondioxide (CO2), argon, krypton, and helium lasers, as well as the lasers with vapors of heavy metals. Liquid lasers use organic dyes dissolved in suitable liquid medium. There are also semiconductor lasers (in the form of light diods, LEDs). Upon transmitting an electric current through them they emit laser light in the infrared and visible area.

Table 1. Classification of lasers used in medicine

Name Active medium Wavelength Light color

Carbon dioxide СО2 10.600 nm infrared Argon Ar 488 and 514 nm blue-green Copper vapor Cu 511 nm green Copper bromide Cu 578 nm Yellow Krypton Kr 521 and 530 nm Green 568 nm Yellow Neodymium- Neodim- yttrium- 1064 nm infrared yttrium aluminum garnet KTP Potassium titanilphosphate 532 nm green

Ruby Chrome 694 nm red Alexandrite 755 nm red Organic dyes Rhodamine 504 nm green fluorescein, coumarin, 400-1.000 nm acridine red (577 and 585 nm - yellow)

CO2 laser emits at 10.6 nm and has low penetration into tissues, due to the strong absorption of its beams by water. There is no selective absorption of its light by the tissue pigments. The CO2 laser exerts only surface action on tissues, causing evaporation when unfocussed and cuttings when focused.

Unfocused CO2 laser is used to vaporise the surface layer of skin in cosmetic surgery of the face. Nd- YAG-laser emits at 1.06 nm and its light has a greater penetration into tissues. Therefore, it has a poor ability to evaporate tissues but increased ability to thermally coagulate them in depth. This explains the demonstration of haemostatic effect by the Nd-YAG-laser, and the absence of this effect in CO2 laser. Nd- YAG-laser and CO2 lasers are both used for resection of the tumor tissues, and the former is preferred when the tumor is localized in well vascularized tissue, while the second is used mainly to remove tumors of bone and nervous tissue. In general lasers can replace mechanical augers and drills used in dentristy to remove tooth caries, for shaping teeth and placing fillings, dentures and more. Unlike the mechanical action of drills on teeth, the laser action is painless, sterile and of greater quality.

4. LINEAR ACCELERATORS FOR RADIOTHERAPY OF TUMORS

The ability of ionizing radiation to kill the cells and tissues is used as a modality for the therapy of tumors. This is the so called radiation therapy, which is applied for less deeply rooted tumors as those of brain, head and neck, lung, malignant lymphomas, mammary gland, the male and female genitalia, urinary tract, gastrointestinal tract, skin tumors and thyroid tumors. Radiosensitivity of cells strongly depends on the absorbed radiation dose. However, there are additional factors of great importance including the time interval in which the dose is taken up, the type of cells and especially - the type of radiation. Radiation comprising particles with greater mass and energy is more damaging because it stronger interacts with the molecules of the cells. Therefore alpha-rays are more damaging than beta-rays and, compared to them, the gamma-rays and especially Ro-photons are the least damaging. The above explains why the classical methods for radiotherapy of tumors (exposure to Ro-rays and gamma rays) are lately displaced by irradiation with fluxes of charged particles - electrons, protons, ions which have mass and higher energy. Such fluxes are obtained using the so called linear particle accelerators that are gradually replacing the older Reontgen systems for percutaneous radiotherapy and apparatus for telegammatherapy.

Fig. 1. Schematic diagram of a linear accelerator of charged particles.

The linear particle accelerators are used for initial acceleration of the charged particles emitted by an electron gun or ion source. Unlike the cyclic accelerator the linear accelerator accelerates the particles transmitting them only once through its accelerating unit. Linear accelerators are used in radiotherapy and radiosurgery as a source of high energy electrons (maximum energy of about 25 Mev) or photons (maximum energy about 20 Mv). For comparison, the apparatus for telegammatherapy use radioactive isotope cobalt-60 as a source of gamma-rays having much lower energy of 1.17 and 1.33 Mev. The main advantages of linear accelerators are: A) They emit high-energy particles which provide radiation with greater cytotoxicity and penetration ability allowing it to reach deep-seated tumors. B) Each linear accelerator can produce particles with suitable energy, lower one for tumors located in shallow tissues and greater one, if the tumor is located in deep tissues. This possibility is absent in the classical apparatus for telegammatherapy that emit only photons with a specific energy. C) An additional major advantage of the linear accelerators is the precise targeting of radiation energy to tumor core with minimal damage to healthy tissues placed above and below the tumor. D) Depending on the type of tumor the linear accelerators could be set at generating either high- energy electrons or high-energy photons as they have a different mechanism of cytotoxicity. For example photon irradiation is the gold standard for treating the brain tumors. In radiosurgery the malignant neoplasms are treated by a single exposure to a powerful and highly collimated (focused) photon beam produced by a linear accelerator called gamma knife (with accuracy of 0.5 mm for brain tumors) and cyberknife (with accuracy of about 1 mm for tumors placed on moving tissues). Radiotherapy uses the same technique of radiation, but the radiation dose is divided into multiple fractions each applied after certain periods of time. The basic scheme of linear accelerators was invented by the physicist Leo Szilard in 1928 and contains a source of particles and accelerating device (Fig. 1).

Fig. 2. An outlook of linear accelerator.

The source of particles can be electron gun emitting electrons or ion source emitting protons or heavy nuclei. The source emits particles into separate groups each one having short duration. The initial particle velocity is low, which requires the particles to be accelerated by the electric field produced by sequentially arranged cylindrical electrodes. When a group of particles approaches given electrode a pulse of positive voltage is fed on the electrode creating electric field that attracts and accelerates the particles. When the group passes through the electrode the electrode voltage is reduced to zero, and the accelerating voltage is transmitted to the next electrode. Thus, the particles are accelerated to speeds close to that of light, and their energy reaches to about 30 Mev. The source of the impulsed voltage is an electric generator based on an appropriate vacuum tube (klystron or magnetron), which feeds the electrodes with electromagnetic pulses of very high frequency (up to 3000 MHz). The frequency and amplitude of the pulses may vary, resulting in different acceleration and energy of the particles, respectively. The obtained flux of accelerated particles moves horizontally and before being directed to the patient it is deviated by 90° with the help of strong electromagnet (Fig. 1). Finally, the flux is used in two ways: 1) it is directed to the patient, or 2) it is guided to a target of hardly-melting metal, wherein a flux of high-energy photons is generated. The resulting photons, in turn, are directed to the patient. In the first case, the beam of accelerated particles must pass through a multi-plated collimator before being directed to the therapeutic area of the patient. Collimator adapts the irradiation flux to the contour of the tumor in order to achieve a maximal therapeutic effect with minimal damage to surrounding healthy tissues. For this purpose the collimator contains a bunch of computer-controlled tungsten plates (up to 160). Each plate is motor-driven and could be inserted to a certain depth into the flux absorbing a portion of it and remodeling the cross section of the flux. Upon the irradiation of the patient the collimator rotates isocentrically about an axis that passes through the center of the therapeutic area. This ensures the radiation to be concentrated in the therapeutic area while the dose absorbed by surrounding healthy tissue is reduced. Fig. 2 shows the appearance of a linear accelerator for radiation therapy of cancer. Prior to radiation therapy, patients undergo a CT scan and sometimes also an MRI scan in order the team of specialists to obtain more details about the therapeutic zone. The practical application of radiotherapy is divided into four stages: 1) planning of dose according to the type of therapeutic area, 2) review of the treatment plan, 3) check and re-planning of the procedure and 4) performing radiation therapy. A systematic review of reported errors and incidents indicates most problems are caused by the medical personal itself. They are chiefly due to incorrect positioning of the collimator, resulting in damage to the surrounding healthy tissues. The type and energy of radiation are chosen depending on the type of tumor and the depth of its localization. For example, the level of energy can be 7 MeV, 11 MeV, 15 MeV and 17 MeV, and the radiation may consist of electrons or photons. There are additional opportunities for exposure mode: radiotherapy with modulated intensity (IMRT), three-dimensional conformal radiation therapy (CRT) and stereotactic radiation therapy. Stereotactic radiotherapy is very precise radiation therapy, in which the tumor is irradiated with very high doses in one or more sessions of exposure. The usual dose of radiation, which is measured in monitor units (MU), is between 50 and 500 per minute. Stereotactic radiotherapy allows doses of up to 2000 MU per minute, reducing the duration of the treatment to 60 minutes or even less. Normal workload of a linear accelerator is about 450 patients a year. In 2013 Bulgaria had 5 linear accelerators and the building of another 14 linear accelerators is planed. The average price of each accelerator is 3-4 million euros, half of which is the cost of installation. The accelerators are installed in a bunker at a depth of 20 meters below ground. The walls of the bunkers are 150 cm thick and have a special lead protection which protects staff and other patients. The whole building has an area of over 500 m2, including a waiting room, changing room, procedure rooms and offices.

5. PHYSICAL BASIS OF MAGNETIC RESONANCE IMAGING (MRI)

This type of imaging is based on the so called nuclear magnetic resonance of the atomic nuclei, hence it is also called nuclear magnetic resonance computed tomography, or MRI for short. For this purpose, the patient's body is placed in a strong permanent magnetic field and simultaneously irradiated by radio electromagnetic waves from the MHz region. The nuclear magnetic resonance occurs when the atomic nuclei, possessing magnetic moments, absorb the energy of the radio waves with a specific frequency (resonance). The absorbing nuclei in turn emit similar radio waves, which are detected. Most often measured is the resonance of the nuclei of the hydrogen atoms in water molecules, which are the most numerous. The nucleus of the hydrogen atom is a single proton, which has a strong magnetic moment interacting with the external magnetic field. For different tissues, the specific emission of these nuclei (protons) is proportional to their concentration. Thus, based on computer analysis of proton concentration a three-dimensional image of tissues and organs could be constructed. In many properties this image is superior to the image obtained by X-ray computed tomography. The main advantage of MRI imaging is that the patient does not receive any dose of ionizing radiation and, in general, the usual contrast agents are not needed. Injecting suitable isotopes such as 31P and 19F in the patient the resonance of their nuclei could be determined allowing the distribution of the corresponding bioactive compound in certain organs to be examined. For example, by determining the resonance of 31P nuclei the active (excited) zones of the cerebral cortex can be observed. The permanent magnetic field is produced by a permanent magnet, strong electromagnet or superconducting electromagnet. Permanent magnet assembly does not consume electrical energy and does not require refrigeration. However, it creates a weak magnetic field and is applies mainly in MRI systems for the pre-operative examination. The superconducting electromagnets are most often used in MRI systems. They contain coils, made of niobium-titanium alloy, placed in a cryostat containing liquid helium (-271.3 °C) or liquid nitrogen (-196 °C). At such low temperature the so called state of superconductivity (very low electrical resistance) sets in the wires of the coil. This type of electromagnets creates a powerful magnetic field at a minimal cost of electric energy, although it requires significant amounts of liquid helium or liquid nitrogen. Nuclear magnetic resonance occurs under the following two conditions. First, the patient's body should be placed in a strong permanent magnetic field. This field orients all nuclei having magnetic moments (e.g. protons of the hydrogen atoms of water) in a same direction and these nuclei start to rotate around the axis of the magnetic field, i.e., to make precession. The rotating nuclei have a same frequency and different phases. The imposed orientation of the nuclei creates an internal magnetization within the tissues, i.e, a longitudinal magnetic field parallel to the outside field and a small transverse magnetic field perpendicular to the outside field. In addition, the nuclei making precession emit their own radio waves. Second, simultaneously with the permanent field the body of the patient must be irradiated with radio waves. When the frequency of irradiating field becomes equal to the frequency of precession of the nuclei a resonance occurs. As a consequence the nuclei, making precession, begin to rotate at a same phase and to absorb the energy of radio waves. The resonance changes the precession of nuclei reducing the longitudinal and increasing the transversal magnetization of the tissue. Upon stopping the radio wave irradiation the nuclei of examined tissue relax, i.e., they restore their normal precession. This transition of the nuclei precession from "disturbed" into "relaxed" state consumes time described by the relaxation times T1 and T2. T1 is the longitudinal relaxation time of atomic nuclei, which for normal tissues lies in the range from 500 to 1000 ms. T2 is the transverse relaxation time and for normal tissues has the values from 50 to 100 ms. T1 is associated with the energy the nuclei lose towards their environment and T2 corresponds to the exchange of energy between the mere precession nuclei. The relaxation times depend on the molecular structure and physical state of the tissue. Nuclear magnetic resonance was discovered at the early decades of XX century. More than 40 years it has been used in radiospectroscopy. For medical purposes it was applies as early as in 1971, when Raymond Damadian proved that there is a difference in the relaxation times of normal tissues and their tumor counterparts. In 1976 Moor and Hinsaw obtained the first images of human tissues by magnetic resonance tomography. The first NMR imaging apparatus was fabricated in 1978. The radio signals, emitted by the precession nuclei of tissues are captured with the help of radio frequency antennas. The antennas are divided into two types: volume and surface ones. Volume antennas are both emitters and receivers of the MRI signals while surface antennas are only receivers. The detected by antennas signals are transferred into mathematical numerical values further collected and processed by a computer allowing the construction of the image on monitor screen. The images of various tissue slices may be obtained on the basis of the measured values of T1 or T2 relaxation times. When the images of the sagittal and transversal slices of the brain are obtained using the data for T1, the white matter has white color (the shortest T1), the gray matter has gray color and the cerebrospinal fluid appears black (the longest T1). In case the image is based on the data of T2 the colors are inverted: the white matter of the brain appears dark (shorter T2), the gray matter of the brain has whitish color between white and gray, and the cerebrospinal fluid appears white (the long T2). To enhance the diagnostic value of the observation special contrast agents are used whose effect is to shorten the relaxation times T1 and T2. There are two basic types of T1 and T2-contrast agents. The T1-contrast agents contain unpaired electrons, i.e. they are paramagnetic compounds. Therefore, they induce a local magnetic field, which reduces the relaxation time T1 increasing the contrast in the T1 mode. The most frequent T1-contrast agent is gadolinium (Gd) - an alkali metal which has 7 unpaired electrons and has pronounced paramagnetic effect. The T2 contrast agents have much stronger magnetic (ferromagnetic, super paramagnetic) properties and are able to induce inhomogeneities in the magnetic field. They reduce the T2 - relaxation time and enhance the image contrast in T2 mode. The magnetite is agent of this kind, administered intravenously in the study of the liver and spleen.

Schematics of a MRI scanner.

The time spent for data collection and change of the regime from T1 to T2 - mode is very small in modern MRI systems. This shortens the study, improves the patient comfort, and inhibits the artefacts related to the motion of patient. This can also suppress the effect of the surrounding tissues (fat, liquids etc.) allowing the differentiation between the solid and cystic structures, blood, necrotic tissues, edema. The main advantage of the method is its ability to present the image in sagittal, coronal and transversal plains and to determine the structure of pathological foci based on the different signal characteristics. Second advantage is the ability of MRI to visualize the large blood vessels in the native studies. For this aim MRI tomography uses the magnetic properties of blood, related to the oxygen binding protein, hemoglobin. Similar to most biological substances the oxyhemoglobin is diamagnetic, while deoxyhemoglobin is paramagnetic displaying stronger magnetic properties. Hence, the oxyhemoglobin is used as a natural contrast agent of blood and this effect is known as BOLD (Blood Oxygenation Level Dependence). The BOLD effect is applied in the NMR-angiography, which makes it possible to obtain images of blood vessels without the use of outside contrast agents. The latter technique allows the finding of vascular constrictions and poor circulation in various organs (brain, kidneys, limbs). Upon stimulation of the motor and sensory centers of brain the local blood flow increases and the concentration of deoxyhemoglobin decreases. This type of testing the brain functions is known as functional MRI method. A third advantage is the ability of NMR scanner to synchronize the data collection with the heart activity. The images produced by MRI systems can be further processed in order to increase the contrast, measure distances, obtain histograms and reconstructions. Images can be transferred to television screen, printed, archived and copied onto optical or electronic media. Magnetic resonance tomography has a clear advantage in the evaluation of pathological processes in large vessels, chest wall, mediastinum, muscle and skeletal system, CNS tumors, multiple sclerosis, spinal injuries, tumors in the abdomen and pelvis. Based on MRI imaging several diagnostic types are possible like diffusion, perfusion and functional magnetic resonance diagnostics. These are still tools for science research, but they increasingly find use also in routine clinical practice. Functional magnetic resonance diagnostics can be applied in psychiatry and studies of memory. There are, however, absolute and relative contraindications for the application of magnetic resonance tomography: • the presence of pacemaker is absolute contraindication; • the list of relative contraindications include the presence of metal implants (depending on the nature of the implant and implant site), foreign bodies, the first three months of pregnancy, claustrophobia.

6. R ADIONUCLIDE METHODS OF DIAGNOSIS: POSITRON- EMISSION TOMOGRAPHY.

Modern medicine uses mainly the following radionuclide diagnostic methods: positron emission tomography (PET), single photon emission tomography (SPET) and scintillation (gamma) camera. Positron Emission Tomography ( PET - CT) measures the concentration of a tracer compound labeled with a positron emitting radionuclide, injected in the tested patient. This diagnostics uses tracer compounds which are involved in the metabolic processes and mimic the behavior of natural sugars, proteins and oxygen in the body. Injected into the patient, the radiotracer distributes in various tissues and organs. The radionuclide in the tracer compound decays and emits positrons which subsequently annihilate on contact with electrons of the body and generate photons. The scanner detects these photons and constructs an image corresponding to the biological activity of cells. Positron emission tomography apparatus is known as PET - scanner or PET - camera.

Fig. 1. The basic unit of a PET - scanner.

The classical methods for imaging (ultrasonography, radiography, X-ray and MRI - CT) depict the structure of internal organs and its changes at various diseases. Typically, the dangerous diseases change the structure of organs and tissues slightly or at later stage. By contrast, positron emission tomography depicts the physiological exchange of certain metabolites in certain organ. At a state of disorder and illness the metabolic processes change at an early stage of the disease when there are no changes in the structure of organs and tissues. Thus, PET-scanner detects alteriation of the function of organs and helps the early detection of pathological conditions long before the appearance of pathological changes in the structure. This is very important in the diagnostics of tumors and their metastasis, inflammation, myocardial damage, epilepsy and many other diseases. The PET diagnostic technique is based on physical phenomena related to the nuclear and particle physics. Immediately prior to testing, proper trace compound, labeled with β+-radionuclide is injected in the patient's blood. After a short time, to allow the equilibrium distribution of the radiotracer in tissues, the patient is placed in the center of a large hoop. A great number of photon detectors are mounted on the hoop (Fig. 1), each detector containing a scintillator and a photo-electronic multiplier. Each pair of opposite detectors is connected to a common electronic amplifier, which amplifies only pairs of signals induced by photons having the same energy of 511 kev and arriving in both detectors simultaneously. This combination of photodetectors and common amplifier operating in such a manner is called coincidence circuitry. Next, the mere scan is conducted in order to measure the distribution of activity (concentration) of the radiotracer in the tested tissues and organs. The radionuclide emits positrons, which are the anti- particles of electrons. After travelling a short distance (about 1 – 2 mm) the positron strongly reduces its speed and ends his way at the first confronted electron. The positron and electron both annihilate (disappear), giving birth to two photons. The most important advantage of the method results from the fact that these photons have the same energy of 511 kev and are emitted simultaneously and in opposite directions at an angle of 180° ± 0.4°. Subsequently, these photons arrive at the same time into a particular pair of opposite detectors. The electric signals produced in these detectors will be simultaneous, will have the same, specific energy and will be enhanced by the coincidence circuitry (Fig. 1). As the coincidence circuitry amplifies only signals entering into strict temporal and amplitude windows, other signal due to photons, scattered from other parts of the tissue and having other energies are not transmitted. If two opposite detectors generate two simultaneous signals (this is the so-called "coincidence event"), it can be concluded that the place of annihilation lies at the line joining the two detectors. This allows the PET-scanner to use the so called electronic collimation of photons, which is superior to the mechanical collimation used by other diagnostic methods. This increases the sensitivity and spatial resolution. While in other methods of nuclear diagnostics (SPECT, gamma-camera, scintigraphy) the spatial resolution is 12-20 mm, in modern PET scanners it is 2-3 mm. The following radionuclides are used in PET scanners: 18F (with a half-life of 109.8 minutes), 15O (2.03 minutes), 13N (9.97 minutes), 11C (20.3 minutes) and 68Ga (67.8 minutes). The maximum positron energy of the emitted positrons is about 1-2 Mev and their mean free path length in water is about 1 – 2 mm. These isotopes have a very short half-life and are obtained by fusion in a cyclotron or in isotope generator near the laboratory, so the method is still expensive and not widely spread. After the nuclear synthesis a chemical synthesis follows aimed at inclusion of radionuclide into suitable compound – glucose, water, ammonia, substances whose molecules bind to specific receptors in cells and others.

Fig. 2. A scanner for positron emission tomography.

Most frequently (95% of cases) the PET scanners use 18F, incorporated in the molecule of 2- fluoro-2-deoxy-D-glucose (FDG). FDG is a glucose analogue, which is taken up by cells and is phosphorylated by the mitochondrial enzyme hexokinase. Almost all tissues, except the kidney tissue are unable to remove the highly charged phosphate group added by hexokinase. Therefore, after the take up FDG remains trapped accumulating within cells, where its tracer radionuclide continuously decays. Thus, those tissues - brain, liver and most tumors that require and absorb large amounts of glucose become radiolabeled. This allows the PET-scanner to detect tumors, such as Hodgkin's limphoma, lung cancer, metastatic and relapsed tumors, etc. Prior to the advent of PET scanners the changes in brain could be studied only through postmortem autopsy or by using animal studies. Now, applying radio-labeled glucose the PET scanner can display the temporary course of glucose accumulation in various zones of working brain, i.e. the uptake of glucose in response to the activation of brain via sound, light, movement of limbs and the like (Fig. 3). With the help of other radiotracers the physician can see the flow of blood in blood vessels of different organs, the entry of oxygen into tissues, the advance of several metabolic and immunological processes. Fig. 2. PET- scans of human brain. Shown are the zones of glucose accumulation in response to sensory activation of brain.

PET-scanner is used in many areas of medicine, particularly in oncology, endocrinology, neurology and psychiatry, for the early diagnosis of malignant tumors, screening of coronary disease, for non-invasive examination of brain function, for the diagnosis of mental conditions, schizophrenia, epilepsy, for study of Alzheimer's, Parkinson's and Down's syndromes, for localization of epileptic foci, and the like. The scan duration is 30 to 90 minutes. By scanning the entire body the physician could establish where is the primary tumor, is there a capsule and metastasis to other organs and systems. The PET- scanner can show the exact location of metastasis and how the treatment affects tumor cells. PET scanning makes it possible to show the cellular metabolism of tumor cells and their level of activity and malignancy. At present the PET - scanner is the only means for diagnosis of tumors in their initial, "zero" phase. PET - scanner can show the viability of the heart muscle and give hints to the cardiac surgeons about the type of heart surgery and the number of bypasses needed. PET - scan exposes the patient to the effects of ionizing radiation, whereat a single dose of about 5-7 mSv is received. Almost the same dose ionizing radiation of 6.5 - 8 mSv is received with a single CT scan of the lung. Frequently, modern diagnostic practice uses the combination of PET / CT scans in which the dose rises to about 23-26 mSv. For comparison, a single X-ray scan of the lungs results in much lower dose of about 0.02 mSv. The average annual dose received by an individual from the natural background radiation is about 2.2 mSv and for pilots it is about 4 to 9 mSv.

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