Transcription
Hello everybody and welcome back. Today, we're going to be looking at the hardware of the CT machine itself, and to do so, we're going to virtually build our own CT machine here.
Now, we've seen this diagram before where the patient lies on a bed within the bore of the CT scanner, and lasers ensure that that patient is dead center within the bore. We've got the housing of the CT scanner. That when we remove that housing, we expose the inner components of the CT machine.
Now, we're going to look at each and every one of these components today and discuss why those components are needed in order to generate a high-quality CT image. We'll look at x-ray generation, we'll look at x-ray detection, and we'll look at x-ray beam manipulation by filters, by collimators, and by the antiscatter grid. We'll also look at the mechanism that allows us for rotation of the CT machine without getting all our wires caught up as this heating machine rotates.
So, let's start by looking at x-ray production. Now, when we look at x-ray production, we look at two specific electrodes that sit within the X-ray tube: the cathode and the anode. We'll start by having a look at the cathode. This is a yellow structure here. If we zoom in on the end of the cathode here, we can see that machined into the cathode face is what's known as a filament. This filament is made of tungsten. It's got a high atomic number, it's really tolerant to a lot of heat, and what we do is we run electricity, we run current through that filament, and the filament gets extremely hot. And that heat within the filament allows electrons within the tungsten to go and be liberated to the surface of that filament. So, as we run current through the filament, we can get electrons accumulating at the surface of the filament.
We can look at this cathode in cross-section. I've drawn a schematic diagram here. We've cut through the filament here, and this cup here is what's known as the focusing cup. At this point, we've got electrical current running to the filament itself, the tungsten filament, and we've got a current running to the cathode, the cup of the cathode here, the focusing cup. The cathode itself is a negatively charged electrode. It's important to remember that at this very moment, we've got electrons on the surface of the filament here.
Now, what we need to do is get those electrons that are sitting on the surface of the filament and accelerate them towards what's known as the anode of the X-ray tube. The anode here, represented in blue. Now, the anode, relative to the cathode, is relatively positively charged. Now, we know that electrons have a negative charge, and because that anode is relatively positively charged, we've got a charge differential between the cathode and the anode itself. And that charge differential is going to allow electrons to accelerate between the cathode and the anode, from the cathode towards the anode here. This red is representing those electrons being accelerated across.
This is now what's known as the tube current. The number of electrons that are available at the surface of the filament is proportional to the current that we're passing through that filament. The more current we pass through the filament, the more electrons are available, and ultimately, the more electrons we're going to get within this electron beam here. The tube differential, the kilovolt potential or the kilovolt peak between the cathode and the anode, is going to determine the energy at which those electrons are accelerated across this gap here. It's also going to influence the number of electrons crossing this gap here. So, the number of electrons is influenced by the filament current, as well as the tube potential, and the energy of these electrons is only due to the KVP, the peak tube voltage across the cathode and the anode here.
Now, these electrons are striking the anode, and that point at which they strike is what's known as the focal spot. The focal spot is where x-rays are actually produced. Now, if we were to accelerate electrons randomly from the cathode to the anode, we would get those electrons spreading out. What we can do is we can actually focus those electrons down onto a specific point known as the actual focal spot, and in order to do that, we use the focusing cup that we mentioned. Now, the cathode is negatively charged, and electrons are negatively charged. So, as electrons are leaving the filament and heading towards the anode, we're going to get this negative charge from the cathode repelling some of those electrons, and that repulsion of those electrons is going to focus the electrons down onto the focal spot.
This is what's known as an unbiased focusing cup, where the current to the filament and the current to the cathode are linked. If we unlink those or link those, we get what's known as a biased focusing cup, where we can control the degree of focusing that's produced by the focusing cup here.
So, now we've got electrons hitting the anode at the actual focal spot. We can do a diagrammatic representation of our anode here. We're looking at the anode head-on here, the face of the anode. Now, when electrons hit the anode, we say that x-rays are produced at that point. Only 1% of those interactions, or less than 1% of those interactions, are actually going to produce x-rays. More than 99% of those interactions are collisional or thermal interactions that are going to produce heat at the anode. That's a major factor that the anode has to deal with: all this production of heat at the actual focal spot. And one way in which the anode deals with all that heat is by rotating. That's why we've got this circular anode here, and you can see this dotted track that I've drawn on the anode here. As the anode rotates, the actual focal spot moves around the anode and dissipates heat over a larger surface area.
The anode is also made of tungsten, one because of its high atomic number, and we're going to see why that's important when we create x-rays, but two, that tungsten has a very high melting point, it tolerates heat well, and we can make an alloy of the surface of the anode here by adding 10% rhenium to the surface. So, we've got a rhenium-tungsten alloy that prevents cracking of that surface. We also have cooling oil that dissipates heat away. Many mechanisms that allow the anode to deal with all that heat.
We also notice that the anode has an angle here. There's a beveled edge to the anode. If we look at that angle side-on now, now we're looking at the edge of the anode, we can see that that angle here is what's known as the anode angle, and where electrons actually hit the anode is the actual focal spot. And based on the anode angle, we're going to get what's called an effective focal spot. The effective focal spot as it heads towards the patient itself, and changing that anode angle is going to change the effective focal spot. The narrower or the smaller the anode angle, the smaller the effective focal spot, which is going to ultimately give us better resolution, spatial resolution.
This is really important when we look at x-rays because a single x-ray is covering a large region of anatomy, and we need a big x-ray field size in order to cover all of that anatomy. Less so in CT, and as we'll see, our CT beam is very narrow. We're just looking at single slices of the patient. We can afford to have very small anode angles because we don't need that large field of view, and the change in the effective focal spot as we move through that field of view is going to be less marked as it is in a normal conventional radiograph.
Now, if any of these concepts are foreign to you, I'd encourage you to go and look at my X-ray physics learning pathway. We cover all of these in much more depth. This equation here that helps us determine the effective focal spot size is what's known as the line focus principle.
Okay, so now we've taken electrons, we've accelerated them towards the anode, and they collide with the anode. Those are called bombarding electrons, and we say that heat is produced predominantly, but also x-rays are produced. And they're two separate interactions that can produce x-ray radiation. Remember, x-ray radiation is just electromagnetic radiation at a certain frequency.
So, we've got our X-ray tube here. This is a bit more of an advanced diagram where we have our cathode, the electrons accelerating towards the anode. You can see that that anode is rotating. These rotors or stators or motors allow for the anode to rotate rapidly. You can see that effective focal track of the focal spot here, dissipating heat over a larger surface area. This is all within a glass housing that provides a vacuum, so those electrons that are accelerating towards the anode aren't interacting with other atoms. And then we've got this all sitting within this oil, this cooling oil, that allows for convection of heat away from the system. The casing of the X-ray tube here is made of a material that x-rays can't penetrate, lead for example. You see these x-rays represented by the red lines are not able to exit the X-ray tube. Only a small window allows x-rays to exit the X-ray tube. This is called the window of the X-ray tube.
We can also see we've got power supply to the cathode and to the anode. We rectify that current, we smooth out that current to make sure that we've got a smooth potential across this gap here, and we don't have alternating current where the electrons pulse with differing energies. So, we've got x-rays leaving the X-ray tube through the tube window here. They're now heading out towards our patient. We can see that those x-rays head out in a beam like this. This is our x-ray beam that's heading towards the patient, and in a future talk, we're going to look specifically at beam geometry and why a lot of these angles become extremely important.
Now, I said we can generate x-rays in two different ways. The first way we can generate x-rays is what's known as Bremsstrahlung radiation. We have our atoms within the anode here. This is a tungsten atom. You can see the tungsten has 74 electrons within it. It's a large atom, it's a heavy atom, and we've got an electron that's coming from the cathode and it's being accelerated towards this tungsten atom at an energy that's proportional to the tube potential, the KVP. That electron is going to accelerate towards the atom, an interaction is going to occur, that interaction is going to produce x-ray radiation called Bremsstrahlung radiation, and it's this x-ray that's heading out towards our patient.
So, what happens? This electron is being accelerated towards the tungsten atom, and the electron is negatively charged. It experiences an coulombic force, an electrostatic force, an attractive force between itself and the positively charged nucleus of the tungsten atom. Depending on how close this electron is to the nucleus, we are going to get an attractive force that's going to slow down that electron. That electron is going to lose kinetic energy and deviate slightly. That loss in kinetic energy needs to be accounted for in a closed system. We know that energy is conserved. The loss in kinetic energy is released in the form of electromagnetic radiation, an x-ray. So, the degree of kinetic energy loss is proportional to the energy of the x-ray that is released. That's really important.
The attractive force between the electron and the positively charged nucleus is inversely proportional to the square of the distance between that electron and the nucleus itself. So, the further away the electron is from the nucleus, the weaker that attractive force is, the weaker that coulombic force is, the less loss of kinetic energy, so the less slowing down there is of this electron, and we get a much weaker x-ray released. The closer that is to the nucleus, the higher that attractive force, the more slowing down we get, and the higher the energy of the x-ray that we're releasing is. This is what's called Bremsstrahlung radiation.
I'm going to draw this slightly differently to show you the spectrum of x-rays that we release when we get Bremsstrahlung radiation. Imagine the cathode being this side, we're accelerating electrons towards the target. We've got one tungsten atom here. We're accelerating them at an energy of 100 kiloelectron volts (keV). The number of electrons here is determined by the filament current, as well as the kilovolt peak, the potential that's created between the cathode and the anode. Those electrons that directly hit or almost directly hit the nucleus are going to lose all of their kinetic energy, and that loss of energy is then going to be released in the form of an x-ray here, a Bremsstrahlung radiation. That energy is going to be equal to the max energy of that electron, which is going to be 100 kiloelectron volts. The further away from the nucleus, the lower where those x-ray energies are going to be.
Now, the likelihood of this Bremsstrahlung radiation to occur increases with the radius or the distance away from the nucleus. The further away from the nucleus we are, the more likely this interaction is going to occur. That's just the function of the circumference that we're creating. When we increase the distance away from the nucleus, we know that circumference is 2πr. As we increase or we're increasing the size, we're increasing the area for these interactions to occur. That's why we get a lot more photons, x-ray photons, in the lower energy ranges because these interactions are much more likely. There's more surface area for this to occur, and we get a spectrum of x-ray energies from zero all the way to our KVP, which is 100 in this case. That's why this is called an x-ray spectrum. This drop-off is a linear relationship, an inverse linear relationship.
So, these x-rays are all heading to our patient. They've got multiple different energies. We call it a polychromatic beam. Now, when we go about actually calculating the mathematics in computer tomography and we generate an image, we assume that there's a monochromatic or a monoenergetic beam, and that's actually not the case. It's important to remember that we need to manipulate this x-ray spectrum to get it as monochromatic as we can, and we're going to see how we do that later.
Now, I mentioned that there's a second interaction that can occur, and this is what's called characteristic radiation. Characteristic radiation produces x-rays of known and predictable energies that are characteristic to the target material. If we change the target material, we're going to get different characteristic radiation. What happens here is our incident electron collides with an inner shell electron of the tungsten atom here. If the incident energy of that electron that's heading towards the tungsten is higher than the binding energy of this, in this example, the K shell electron, it's going to eject that K shell electron. So, we've ionized this tungsten atom, and we've rebounded or reflected our initial electron, our incident electron. We've now created a vacancy within the K shell.
What happens is an L shell electron or even an M shell electron is going to drop down and fill that vacancy. We're going to move from a higher energy state to a lower energy state. Now, remember, energy in a system is conserved. If we go from a higher energy state to a lower energy state, we're going to compensate for that energy loss with the release of electromagnetic radiation in the form of an x-ray. So, you see that dropping down from an L shell to the K shell, and we get a release of characteristic radiation. Now, it's characteristic to tungsten because we can actually calculate the energy of this x-ray radiation based on the binding energy difference between the L shell and the K shell.
Again, I'm going to show you a similar diagram to what we looked at previously, but now I'm just showing you a K, L, and M shell of this tungsten atom. Same electrons are heading towards the target, but now we expel one of the K shell electrons, and either an L shell or an M shell jumps down into that K shell, and x-rays are released based on the differences in the binding energies here. If an L shell is going to fill a vacancy made in the K shell, we're going to get x-ray radiation in the high 50s here. If an M shell was to fill it, then we're going to get x-ray radiation in the high 60s here. This radiation coming here when an L shell electron fills a K shell is what's known as K-alpha radiation, and when an M shell fills a K shell, that's what's known as K-beta radiation. So, we've got K-alpha and K-beta characteristic radiation here.
Now, this happens in the other shells, but the energy difference in the binding energy is very small. If an M shell fills an L shell, we're looking at kind of 8 to 10 kiloelectron volts, going to be very low down on the spectrum here. And if further shells, an N shell, then fills an M shell, we're dealing with less than one kiloelectron volt. Those x-rays are never going to make it through the patient. They're unlikely to even pass through the window themselves because they've got such low energy.
So, we've looked at characteristic radiation and Bremsstrahlung radiation. If we combine them, then we get our unfiltered x-ray spectrum. These are the energies, the range of energies of the x-rays that are heading towards our patient, as well as the absolute photon numbers, the number of x-rays that are heading towards the patient. And that's why it's called a spectrum because of that range.
Now, we're going to look at a concept called filtration later on, where x-rays are attenuated as they pass through different structures within the X-ray machine. If x-rays are leaving the window, we get what's known as inherent filtration. The lower energy x-rays aren't even going to make it through the X-ray window. We can also add sheets of metal, and why that's known as added filtration, when we can further manipulate this x-ray spectrum. We're going to look at that in some detail later.
So, now we've got x-rays in a spectrum heading out towards the patient. They've got multiple different energies, and they're all within one geometric plane known as the X-ray beam here. We need a way of detecting those x-rays. We need a way of figuring out how many of those x-rays have been attenuated by different structures and how many have passed through the patient and been transmitted. The crux of radiography, of computed tomography, of x-rays, is figuring out attenuation differentials between different structures in the body. We need x-rays to be transmitted through the patient, and we need x-rays to be attenuated by the patient, and we need to be able to calculate that difference and then represent that difference graphically on the screen.
Now, in order to do that calculation, we need detectors. Here, I'm placing a row of detectors here. These are what's known as x-ray detectors. They're detecting x-rays. Here, I've got a multi-detector array here. We haven't just got a single detector; we've got many detectors here, and we've got multiple rows of detectors. These are all part of a third-generation CT scanner, and we're going to look at generations of CT scanners later on. So, we're detecting x-rays here. We need some way of digitizing the x-rays that we detect and ultimately processing that digital data and displaying it as an image.
Now, the way we detect x-rays is either by indirect mechanisms or direct mechanisms. I'm going to be quite brief here, but I have looked at this in some detail during the X-ray learning pathway. In indirect systems, is what we predominantly use now currently in CT imaging. What happens is we get an incident x-ray that's converted to light by a layer that's known as a scintillator layer. X-ray radiation is converted to light. That's very important. This layer can be gadolinium oxy, can be multiple different compounds depending on the manufacturer of the CT machine. That x-ray is then converted to light, and that light is then received by a photodiode layer. That photodiode layer takes the light energy and converts it into an electrical current, electrical energy. And the number of light photons is proportional to the number of x-ray photons, and the degree of electrical current that we generate is also proportional to the number of light photons that we have. So, we can see how we've converted x-ray energy into light into an electrical signal. That electrical signal is an analog signal that we need to then amplify and digitize using an analog-to-digital converter, and then we're going to take that digital data and process it later on. We'll look at those when we look at data acquisition and storage and processing of data.
Now, there are also direct mechanisms where we take x-ray radiation and we convert it directly into electrical current. We skip out this scintillating layer, we skip out the light step. In the past, there was a gas chamber known as a Xenon ionizing gas chamber, where incident x-rays created electron pairs, a negatively charged electron that moved towards one end of the gas chamber and its positive ion pair that moved to the other side and generated a current. And then that current again was stored and processed using this element here, which is a common element between these two systems known as the detector element or the DEL. These are individual detector elements here.
Now, when we look at this scintillator layer, this is what's known as a solid-state device, and it's the most commonly used detection mechanism in third-generation CT scanners, which we're using now. These are mechanically scored. They're scraped between the detector elements, and those defects that are created in the surface of this x-ray detector are filled with an opaque filler. It prevents light from scattering out towards other detectors. It keeps that light on one single detector, and it allows us to create discrete data points based on each single DEL within this detector here.
So, now we've created x-rays and we've detected the transmitted x-rays through the patient. Let's place a patient within the X-ray beam here and notice now that x-rays in the lateral parts of the beam here, in the Z directions, they're going to pass through the patient and not hit the detector. And we know that x-rays interact with matter. One type of interaction is scatter, where the x-ray's path is changed based on the interaction with the patient. So, we could get x-rays that weren't going to hit the detector; they could scatter and hit the detector. The detector is going to assume that x-ray came from a straight line, and we're going to get a loss in resolution in our image here because of the scatter that's occurring. So, that's one issue. We've got scatter that's going to hit our detector that wasn't originally lined up with our detector.
The second problem is we've got x-ray radiation passing through a region of the patient's body that's not useful to us because it's not heading towards a detector and it's exposing the patient to ionizing radiation. We've got an increased dose that's not providing us with any clinical value. What we need to do now is narrow down that x-ray beam so that the x-ray beam is only incident on the detector. We can place two metal sheets that prevent x-rays from passing through, and they're what's known as collimators. We can place collimators here and we narrow down that electron beam. See what's happening here? As we place the collimation in here, we've now made sure that the beam only hits our detectors. We've reduced the patient dose, but we've also reduced the amount of scatter that's coming from regions of the body outside of the field of view here. So, we've reduced dose and increased image quality. That's why collimation is extremely important in computer tomography and in x-ray imaging.
Okay, so now we've narrowed down that beam to ensure that it just hits our detectors, and we know that these x-rays can interact with matter. So, let's place three different attenuating structures. Pretend these are within our patient. X-rays passing through this patient can either be transmitted through a structure, attenuated by that structure in a process known as the photoelectric effect, or they can be scattered by that structure. We looked at scatter previously, looking at collimation, that from the Z plane. Now, this scatter is happening within the X-Y plane, within our single slice, and this scatter could be Compton scatter or Rayleigh scatter. These are all interactions with matter.
Now, the problem is this detector element that's detecting the scattered x-ray is going to assume that x-ray passed straight through this structure here, which would have normally attenuated that x-ray, and we're going to misregister scatter. So, we've already reduced scatter by collimating the beam. Now, we're going to reduce scatter by adding what's known as an antiscatter grid in here. You can see these cross-hatching grids here, these highly attenuating septa that will prevent x-rays passing through them. You see, each one of these septa will not allow x-rays to pass through them, much like the collimator didn't allow x-rays to pass through them.
Now, the antiscatter grid is aligned perfectly with the incident x-ray beam, so x-rays that pass through in the correct angle are going to pass through this antiscatter grid. X-rays that have been scattered are going to hit the septa prior to reaching our detector; they're going to be attenuated and not reach the detector. So, let's place the detector within our machine here, and you see how we're no longer going to misregister x-rays here.
So, you can see that this antiscatter grid is fairly unique to CT imaging because we've got cross-hatching of those septa. Remember, in x-rays, we generally have them up and down, just in one orientation. Now, there are certain properties of these grid septa that are going to determine how many scattered x-rays are going to be able to make it through this antiscatter grid. We've got what's called the grid ratio. The higher the septa and the narrower the width, or the narrower the interspace material between the septa, the higher the grid ratio is going to be, and the fewer scattered x-rays that are going to pass through this grid septa. The wider the interspace material and the lower the height of the grid, the more scattered x-rays that are going to pass through here. That's just basic geometry with angles of x-rays passing through this antiscatter grid.
We also looked at the grid frequency, how often these grid septa come up, and we're going to, in CT imaging, we're going to match the grid frequency to the width of our detector elements so that the antiscatter grid can lie over those regions that we scored on the detector elements and filled with an opaque filler. So, those antiscatter grids are going to block primary x-rays that weren't even going to hit the photodiode themselves; they were just going to hit that opaque filler on our detector elements.
Now, there are a couple of differences here between this antiscatter grid and the antiscatter grid that we use in conventional radiography. When we are taking a CT scan, let's place our patient within the middle of the CT scan. The geometry between the X-ray source and the detector is fixed. They're lying on the same gantry, and they're rotating together. In third-generation CT scanning, nothing is going to change the angle of the X-ray source to the detector itself, and that allows us to create this fixed angle between the source and the detectors. Our grid is not going to change; it's built into the machine. We're not taking the grid out and putting a different grid in. In x-ray and conventional radiography, we place the antiscatter grid in front of the detector and we then shoot the x-rays through the patient.
Now, two things may occur here. We may want to change the angle of that x-ray. Say, we haven't got the angle exactly right, so we're taking a shoulder x-ray and we want to get that Y view perfect. We may need to change the angle slightly and take the x-rays from a different incident angle. You can see here that the changing in angle, if we were to have a cross-hatched grid, would mean that none of these primary x-rays would actually make it through the antiscatter grid. That's why in conventional radiography, we've got those grids that just have straight lines up and down. The patient also might be lying on a bed; we might be taking a mobile x-ray, and that bed angle is not perfectly fixed like we have in computed tomography.
Okay, so now we've built most of our machine here. We've got x-rays that are heading through the patient towards our detector, and I said to you that we've got a spectrum of energies within those x-rays, and we want to manipulate those x-rays to make sure they're more monoenergetic than this polyenergetic beam here. What we can do is add what's known as filtration. We can add a metal sheet here before the x-rays reach the patient, that's going to manipulate the x-ray spectrum. So, if we add that sheet in here, look what's happened to the x-ray spectrum. Let me do that again. Add the sheet in there. What's happened is we're preferentially attenuating the lower energy x-rays. Only the higher energy x-rays are making it through this filter towards the patient. This filtration has reduced the total number of x-rays heading towards the patient, but it's increased the average energy of those x-rays that are heading towards the patient. We've reduced the beam quantity, the number of x-rays, but increased the beam quality, the average energy of those x-rays.
Now, the interactions that occur here between the incident x-rays and the filter is what's known as the photoelectric effect, the same thing that's happening in the body when x-rays are attenuated. And the probability of the photoelectric effect to occur is based on the density of the filter that we put in, as well as the atomic number of the filter that we put in, and it's inversely proportional to the energy of the incident x-rays. The higher the energy of the incident x-ray, the less likely the photoelectric effect is to occur. That's why higher energy x-rays make it through and lower energy x-rays don't.
Now, what effect is going to have on the image? By reducing all of these lower energy x-rays, you see here the photoelectric effect occurred in our filter with these lower energy x-rays at a high rate. The same thing happens in the body. These x-rays are never going to make it through the patient's body and are never going to give us useful information that reaches our detectors. The only thing that these x-rays do is contribute to patient dose. So, adding a filter is going to remove that dose, and it's going to have no difference on the quality of the x-ray that we get at the end of the day here.
We can also get a different type of filter known as a beam-shaping filter. The initial simple filter was just to change the x-ray spectrum to increase the average energy and reduce the number of x-rays. A beam-shaping filter is going to influence the number of x-rays that reach our detector. Let's have a look at it in closer detail.
If we place a patient within the X-ray field here, in the X-ray beam, we can see that the patient shape is not uniform. Say we've got an abdomen here, and if we look at two separate x-rays passing through that patient, they're going to pass through differing lengths of patient based on the orientation of the patient here. The flanks of the patient are thinner, the central part of the patient is much thicker. When we detect x-rays, what we want the difference in that detection, the difference in the number of photons, to represent is the differing attenuation properties of the patient, not the differing thicknesses of the patients. Ultimately, what we're trying to calculate is attenuation differences.
Pretend this is a unique patient, and they uniformly attenuate x-rays at the same rate. Every part of this patient has the same Hounsfield unit value. Let's take those two parts where the x-rays went through the patient, and notice how different the length of those paths are. One's very short, the other's double the length. We can take what's known as a half-value layer. We've looked at this previously. The half-value layer is the thickness of tissue that will reduce the x-ray intensity by half. So, if we were to shoot x-rays at these two lengths here, let's say we've got 200 x-ray photons going to each part here. 200 x-ray photons traveling through one half-value layer of tissue is going to leave us with 100 x-ray photons. 200 x-ray photons traveling through two half-value layers: the first half-value layer is going to halve the number to 100, the second half-value layer is then going to halve that 100 down to 50 x-ray photons. The attenuation of these two tissues, we've said, is exactly the same. There's no attenuation difference; they should have the same Hounsfield unit value in our final image. Yet, we're getting differing x-ray intensities at the detector here, and that difference is just based on the thickness of the tissue.
What a beam-shaping filter can do is try and account for that difference in thickness and allow the differences in detection to be purely based on attenuation of the tissue. Let's take the same example here, but now we place a beam-shaping filter. This filter is thicker on the sides and thinner in the middle. It's called a bow-tie filter, based on its shape. We have the same x-rays, 200 photons, heading to each of the tissues here. These x-rays pass through the filter, and 120 of those x-ray photons are attenuated out prior to reaching the patient on the other x-ray track, because it's only going through a smaller portion of the filter, only 40 of those x-rays are attenuated out. Now, if these 80 x-rays go through one half-value layer, we're only going to get 40 x-ray photons reaching our detector. If these 160 go through two half-value layers, we'll get 80 attenuated initially, and then of the 80 remaining, we'll get a further 40 attenuated; we'll get 40 x-ray photons. Now, we've evened out; we've shaped the beam to make sure that the photons incident on the detector now are equal, considering that these are the same attenuating tissues here. The difference that we see here, if there were any difference, would be based on attenuation differences within the patient. Very important concept when it comes to filtration in CT imaging.
Okay, I know we're taking long. The last problem we need to solve is that the CT machine, the X-ray tube, needs to rotate, and we saw that the X-ray tube had wiring, and we want to continuously rotate this CT machine, sometimes up to five times per second. This X-ray tube is going to rotate around the patient. That wiring is going to get coiled up more and more if we were to continually rotate. I want to show you how we can get around that using a concept known as slip ring technology.
We've now mounted our X-ray tube onto this blue base here, and we've mounted our detectors onto this blue base here. These blue bases are going to rotate around this green part of the CT machine, the CT gantry here. The green part is remaining still, the blue parts are sliding along the green part. Can you see these blue parts here sliding along the green part, which is completely still? The gantry is completely still. Now, if we were to look on the posterior aspect of this gantry, we can see that we've got these rings. They're known as slip rings. They conduct electricity and they provide power to the X-ray tube here. In these little gaps here, where the blue part of the X-ray tube meets those slip rings, we've got little bristles, metallic bristles, that are going to remain in constant contact with the slip rings as the blue part of the machine rotates. The same thing happens both at the X-ray tube part and the detector part, because these remain in constant contact as they slide over the slip rings, we can get a continuous current supply to the X-ray tube. On the detector part, we can receive information from the detectors continuously and send those off for processing. So, we can supply energy here and receive electrical current here, and it's the slip rings that allow us to continually rotate, continually produce x-rays as we are acquiring the data.
Now, we never turn the x-rays on and off, on and off, as we acquire each different angle around the patient. The x-rays are continuously being produced. The rate at which we sample the data from our detector elements will determine how many different samples at different angles we take as that x-ray machine rotates, and we generally take 1,000 to 3,000 different x-ray projections through a single slice of the patient. So, we're rapidly sampling the attenuation data that we're creating here.
To finish off, we then house all of this within the housing of the gantry. We've got the bore of the scanner here, and lasers to ensure that the patient is directly on. Now, the table is able to move up and down to ensure that the patient is central, but it's also to be able to move through the bore of the scanner as that CT machine is rotating. So, as the patient moves through the scanner, we are radiating, we're exposing different regions of the patient to x-ray radiation, and that allows us to get continual data over a region of anatomy.
Now, you may be thinking, if the patient's moving whilst we're acquiring the data, surely that needs to be accounted for later on, and you're absolutely correct. In fact, there's lots of geometry within the heating machine that we're going to cover in the talk after next, and I'll show you there's lots of considerations. Actually, it's not as simple as we may think initially.
In the next talk, I want to look at the generations of CT scanners. How have we come about this setup that we're using now? Where did we come from, and what have we tried, and why is this the best setup for us? That talk is going to be very brief. We're going to look at five different generations. It's a common question that comes up in exams. They love to ask about the different generations of CT scanners, and it's a question that I will be including within the question bank link below if you're studying for a specific exam. So, I'll see you all in that talk. We've got two talks on the CT machine: the generations of the CT scanner and the geometry of the CT scanner. Then we can get on to the more exciting things like Hounsfield units and how we actually go about creating the image, looking at artifacts, and all the really important things for us clinically. So, until that talk, I'll see you all. Goodbye, everybody.