Analysis of bone by micro-CT General information micro-CT analysis ... along with a superficial look at physical ... crossectional area converted to a bone volume estimate. Measured

Analysis of bone by micro-CT

General information

Method note

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Bone micro-CT analysis

Abstract Micro-computed tomography, or micro-CT, is being employed by a rapidly increasing number of laboratories for osteoporosis and other bone research, while the technology is still in its first decade. Bone biologists (and other biologists also) are a significant group of micro-CT users. This article is intended to give users without an extensive background of physics a better understanding of the methodology with its capabilities and limitations. The background of micro-CT bone analysis will be outlined, along with a superficial look at physical principles underlying the technique. There is a brief examination of the methodology of two important types of micro-CT analysis - morphometric and density measurement. Finally there will be a review of some selected applications of micro-CT in published bone research, including both ex vivo and in vivo analysis of bone.

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1. Introduction Micro-computed tomography, or micro-CT, is being employed by a rapidly increasing number of laboratories for osteoporosis and other bone research, while the technology is still in its first decade. Bone biologists (and other biologists also) are a significant group of micro-CT users. This article is intended to give users without an extensive background of physics a better understanding of the methodology with its capabilities and limitations. The background of micro-CT bone analysis will be outlined, along with a superficial look at physical principles underlying the technique. There is a brief examination of the methodology of two important types of micro-CT analysis - morphometric and density measurement. Finally there will be a review of some selected applications of micro-CT in published bone research, including both ex vivo and in vivo analysis of bone.

“Computed tomography” implies taking x-ray projection images of an object from many angles around the object, and mathematically converting this set images into a stack of crossectional image slices, which represents a 3- dimensional image. These projection images are taken incrementally over a total rotation of either 180 or 360 degrees. While the individual projection x-ray images are 2D images, the rotation of the imaged sample relative to the x-ray source and camera allows the precise 3D location of the scanned objects to be calculated (provided the objects are within the camera field of view at all angles of rotation). This is shown schematically in figure 1.

Figure 1. The geometry of CT: the X-ray source located at S and the detector centred at O’ have a fixed position relative to the scanned object, which rotates on a circular orbit around point O, the centre of the field of view. Point O is the centre in an x-y-z coordinate system. At different angles of rotation (θ) the projection images of the points O and P (that is, O’ and P’ on the 2D detector) have different positions. From these changing positions the absolute 3D locations of O and P are calculated by the reconstruction algorithm.

Medical CT scanners have been an important part of hospital medicine since the invention of the technique by Godfrey Hounsfield in 19721. Micro-CT is CT of

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small objects in specialised scanners which have a very high spatial resolution. The micro-CT imaging process is illustrated in figure 2 with the example of a micro-CT scan of a mouse hindlimb in vivo. Figure 2. Multiple projection “shadow” x-ray images of an object from different angles (upper image) - here a live mouse hindlimb - are analysed by the mathematical techniques of back-projection and convolution, to generate a stack of crossectional images. The two example crossections shown (lower image) are from the tibial metaphysis, from the live mouse, from a 20 minute in vivo scan.

1.1. Micro-CT bone morphometry - an extension of bone histomorphometry.

In the early to mid 1990’s developments in x-ray instrumentation technology allowed commercial development and realistic laboratory availability of micro- CT scanners2. The progress of scientific research has been punctuated by key technical developments which are followed by an explosion of fruitful new investigation. Examples of key imaging technologies in the history of science include light microscopy, telescopes, x-ray imaging, electron microscopy, magnetic resonance imaging, atomic force microscopy and so on. Technical developments run alongside purely intellectual developments in determining the course or scientific progress.

In the case of bone biomedical research, the essential fore-runner of micro-CT bone morphometry was bone histomorphometry. This started with the creation of histological methods for the embedding and sectioning of calcified tissue

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without prior decalcification3. These developments were accompanied by new staining methods (e.g. Von Kossa, Goldner’s Trichrome, in vivo fluorochrome labelling) which enhanced the histological study of bone, allowing the distinguishing of mineralised bone from unmineralised newly-formed osteoid, and measurement of the dynamics of bone formation (figure 3; refs. 4, 5, 6, 7, 8). This in turn was accompanied by progress in microscope image analysis technology such as video cameras and partial automation of analysis, plus the rapid development of personal desktop computers and thus the ready availability of significant computing power. All this together allowed the growth of bone histomorphometry, the quantitative study of mineralised bone tissues. Histomorphometry provided a wealth of new information on the structural and metabolic status of bones from experimental animals and of biopsied bone from human patients. Figure 3. One histological stain popular with histomorphometrists is the stain, which allows simultaneous visualisation of osteoid and some cellular structure along with fluorochrome dynamic labels such as (here) calcein, viewed with epifluorescence.

1.1.1. Stereology

In parallel with these technical developments were developments in the field of stereology, the mathematical formulations which allow measurements made in 2-dimensions (2D) on single or multiple microscope sections through a volume of bone to be converted into estimates of 3-dimensional parameters 9. Thus bone perimeter measurements could be converted to an estimate of surface area, and crossectional area converted to a bone volume estimate. Measured perimeter to area ratio provided the basis for estimates of bone trabecular thickness. These conversions involve some quite complex mathematics and are correctly described as estimates, since they have several significant sources of error. (It is

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worth noting that the “Poincaré Conjecture”, a question regarding the derivation of 3D information from 2D sections, is one of the most complex and enduring problems in mathematics!) A major source of error in deriving 3D bone structural parameters from 2D crossectional measurements is the need to assume an architectural model for the structures being studied. For example one can apply a “plate” model and assume that the (trabecular) bone studied is composed only of parallel flat plates, or a “rod” model, assuming the sample is a regular parallel array of cylinders. Real objects of interest such as trabecular bone depart greatly from such idealised structures.

What micro-CT brings to bone research - and indeed to any research field - is the ability non-destructively and directly to obtain a 3-dimensional image of a scanned object. The word “direct” here is important, implying the imaging method is 3D in a true and fundamental sense, providing a 3D model of an object based directly on imaging measurements. The great advantage of this for bone morphometry is the chance to “side-step” all the mathematics of 2D-3D conversions and model-based estimates, and to make direct measurements on a 3D image model10, 11, 12. So this represents a simplification as well as a significant improvement in the capability to obtain information about the 3D architecture of structures such as bone. Furthermore, having a 3D image model makes possible some wholly new architectural measurements of parameters that are completely inaccessible to analysis based on 2D image slices. This will be discussed in more detail later.

With the advent of micro-CT in bone research the term “morphometry” has started to supplant “histomorphometry” because it is no longer the province of histology alone to obtain bone morphometric data. However it is probably correct to define micro-CT bone morphometry as an extension of the discipline of histomorphometry, rather than a wholly new science. Indeed many bone scientists familiar with the technical difficulties, limitations and labour intensiveness of histomorphometry were enthusiastic early adopters of micro- CT technology, recognising the implications and value of direct 3D bone morphometry from non-destructive imaging. However micro-CT technology in its current state cannot fully replace histomorphometry since it provides only structural or architectural information about a bone sample, and cannot image osteoblasts, osteoclasts, resorption lacunae or osteoid seams on the bone surface (at least not over a wide enough region to be useful quantitatively). So continuity of histomorphometry and micro-CT bone morphometry will continue, and it is recommended that this should be reflected in the adoption by practitioners of bone micro-CT analysis of the ASBMR system of symbols and abbreviations for measured parameters developed for histomorphometry13, even for new 3D-only parameters.

A word of caution is appropriate. The very impressive visual images that can be produced from micro-CT datasets, either in virtual 3D or in crossection (e.g. figure 4), can give a new-comer to the methodology an unrealistic sense of what is possible by the technique. To master micro-CT analysis requires a certain level of understanding of the physical principals and factors involved. This will lead to an understanding of how to get the best results from the technique and of the boundaries of what is possible and what is not, as well as the nature of several CT imaging errors or artefacts that need to be minimised or corrected.

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1.1.2. Non-destructive 3D imaging

The non-destructive nature of micro-CT is one of its key advantages. Bone can be analysed by micro-CT ex vivo following post-mortem dissection out of a rat or mouse, or following extraction by biopsy from a human patient14, 15. Following micro-CT analysis the samples can then go on to be processed and analysed by histology or subject to chemical, biochemical or other types of analysis. In vivo micro-CT scanners are also available which allow bone within a living anaesthetised mouse or rat (see figure 2), extending the principle of non- destructive bone analysis still further16. Another recent technical development is a micro-CT scanner suitable for clinical analysis of the human wrist in vivo 17. This short chapter will outline some important physical principles connected with micro-CT imaging, and then give some examples of application of micro-CT to bone and osteoporosis research. The intention is to assist a newcomer to the field, and especially a scientist planning to commence a bone research project involving micro-CT, to understand the methodology. Such a goal is no doubt too ambitious for a single article, but it is hoped that the goal will be at least partly achieved.

Figure 4. Visually impressive 3D images of samples such as bone (mouse femoral trabecular bone), generated by micro-CT, show the power of the methodology but can give an unrealistic sense of what is possible by the technique. A more realistic perspective is gained by learning the essential physical principles of micro-CT imaging.

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2. Physical principles of micro-CT x-ray imaging

2.1. Interactions of x-rays with matter

X-rays are electromagnetic radiation like light but with higher photon energy. Radiation including x-rays can be used to image the internal structure of objects only in the case where some rays stop in the object and others go right through. X-rays of the type produced by typical laboratory equipment will for example provide no information about structure of gasses (due to total transmission) or about a centimetre thick block of lead (due to total absorption). So partial absorption (or partial transmission) of the micro-CT system x-rays is needed for imaging internal microstructure.

If the whole object has the same partial absorption of x-rays, its image will be represented by a uniform shade of grey, and not be very interesting. So a second requirement for meaningful imaging of internal structure by transmission of x- rays through an object is differential absorption, that is, different parts of the object having significantly different x-ray absorption - but each part still having absorption within the appropriate range for image formation. The best results with micro-CT imaging are obtained from objects in which detailed internal structure coincides with contrasting x-ray absorption.

The type of image formed in this way by the differential transmission or absorption of radiation through different parts or phases of an object, is called a projection image. The projection images are the raw material for the tomographically reconstructed 3d crossection images in micro-CT. It is very important to remember that the quality of the reconstructed 3d dataset is directly determined by the quality of the projection image, in terms of contrast, signal-to-noise ratio, etc. So adjusting scan parameters to get the optimum quality projection images is an important part of the micro-CT analysis process.

Figure 5. The relative importance of the three major types of x- or gamma ray interactions are determined by absorber atomic number (Z) and photon energy (hv). The coefficients τ, σ and κ represent the probabilities of photoelectric absorption, Compton scattering and pair-production, respectively18.

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It is helpful to have a superficial look at the physical theory of how x-rays interact with matter. The three most important modes of interaction between x- rays and matter are photoelectric absorption, Compton scattering and pair production. Their occurrence is determined by two factors: the x-ray energy and the atomic number (Z) of the absorbing element, as illustrated in figure 5. Photoelectic absorption is the predominant interaction mode for low energy photons especially where atomic number is high. The important characteristic of photoelectric absorption is that the incident x-ray photon loses its entire energy in the interaction. This makes photoelectric absorption the preferred interaction mode for micro-CT imaging, since it results in “clean” x-ray removal from the incident beam. By contrast Compton scattering removes only a part of the photon’s energy, and the x-ray photons continue on with a different path from the original beam leaving the x-ray source. If these scattered photons are detected by the x-ray camera they cause scattering artefacts in images.

The third mode, pair production, involves a higher energy photon being converted into an electron-positron pair in the vicinity of an absorber atom. (The positron is short-lived being quickly annihilated to produce a pair of photons.) It is mentioned for completeness but is not relevant to micro-CT imaging, since it occurs only at photon energies well over 1 MeV. (Although it is highly relevant to positron emission tomography (PET) molecular imaging).

The practical implication is that micro-CT imaging of samples with higher atomic number and with lower x-ray energies will make photoelectric absorption the predominant mode, and result in superior imaging results. For bone scanned by desktop micro-CT scanners, photoelectric absorption is generally predominant, which is why bone is an ideal sample material for micro-CT imaging. None-the- less, decreasing applied x-ray voltage in micro-CT systems will increase the relative amount of photoelectric absorption and increase contrast in the transmission image. But too low a voltage results in poor transmission (dark image) and an insufficient camera signal - this is a trade-off that must be found for each type of imaged sample. As typical examples, rodent or clinical biopsy bone samples are best imaged with tube voltage at 40-60 kV, while larger samples such as bones from large mammals or where thick cortical bone is present, require 70-100 kV.

Another important characteristic of photoelectric absorption is its very sensitive dependence on the atomic number of the absorbing material. In general,

CONSTANT

n

Z 3.5

(1.) E

Where is the coefficient of photoelectric absorption, Z the atomic number, n is a value from 4-5, and E is the x-ray photon energy. The fact that Z is raised to the power of 4-5 means that photoelectric absorption is acutely sensitive to atomic number. This fact is of central importance to x-ray imaging in micro-CT, and is the basis of the formation of contrast in micro-CT images (and indeed all x-ray images). It results for example in the strong contrast between bone and soft tissue in an animal. Bone is less than 2 times more dense than soft tissue in terms

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of weight, but due to the presence of Ca (Z=20) the x-ray absorption of bone is greater than that of soft tissue by a much larger factor.

The Z-dependency of photoelectric absorption in fact sets the principal criterion for imaging of objects’ internal structure by x-rays, namely that the structural features that one wishes to image must coincide with gradients or differences in mean atomic number of the constituent material or material phases.

A final general comment should be made about x-ray absorption - that it is exponential in nature. This means for example that if a certain thickness of a material removes half of a beam of x-rays, then twice that thickness of the same material will not remove all the x-rays, but instead only three quarters of them. Three times the thickness will remove seven eighths, and so on. X-ray absorption or “attenuation” is probabilistic in nature, depending on the chance or probability of x-rays interacting with atoms as they travel close to them. X-ray attenuation is described by the equation (number 2) below - note that this relates to the simple and ideal situation where the x-rays are of a single energy and travelling in a parallel beam through a uniform solid.

x I I0e

where I and I0 are x-ray beam intensity before and after travelling a distance x through a solid which has an attenuation coefficient of µ (e is the exponential coefficient, about 2.718282).

2.2. Essentials of micro-CT image formation

The usual essential elements of any imaging technique apply to micro-CT, such as contrast, brightness and signal to noise ratio. It is worth carefully following the instructions of your micro-CT system supplier to optimise these properties in the micro-CT bone images that you obtain. We have noted already that contrast can sometimes be improved by lowering source voltage, due to increase in absolute and relative x-ray absorption in different phases of the scanned object.

2.2.1. Reconstruction contrast limits

During reconstruction, the user selects the image contrast limits. The projection images store raw image information on x-ray brightness and attenuation, typically in the form of 16-bit images (with 65,536 grey levels). The reconstructed crossection images for viewing are greyscale images with a number of grey scales between black and white depending on the bit-depth of the reconstructed images (256 levels if images are 8-bit) It is necessary to specify what values of x-ray brightness from the projection images will correspond to the black and white values of the reconstructed images. Generally the lower limit should be at or just above the zero value for raw brightness. (Raising this lower limit can make the relatively noisy ambient and low density part of the image disappear but this does not actually improve the image’s signal to noise ratio, and can harm the accuracy of segmentation at the surfaces of objects where a more complete density gradient is an advantage.) If density

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measurement is planned with the scanned dataset then the lower contrast limit should be zero. Figure 6. The contrast limits for bone should be set appropriately for density measurement. The lower limit should be zero. The upper limit should be at the upper (right hand) end of the spectrum “tail” representing the highest bone CT density value.

To position the upper contrast limit you need to look at the spectrum of x-ray brightness / absorption for a reconstructed crossection (see figure 6). At the upper end of the spectrum of brightness values associated with the scanned object there is a “tail” representing the highest values of density, or x-ray opacity. Position the upper grey limit close to the top of this higher density tail, as shown in figure 6. Note that all values above the selected upper contrast limit will be saturated in the reconstructed images, having a single grey level (black or white) regardless of the actual x-ray absorption value. The same is true of pixels with absorption values below the minimum selected contrast limit.

2.2.2. Noise

Noise is a particularly important issue in micro-CT imaging since the method depends on an x-ray source with a very small spot size - typically 5-8 µm in diameter - and this strictly limits the maximum source power and x-ray flux. Thus the image signal strength is on the low side for micro-CT. Generally you should check that your background x-ray camera signal is about 40-60% of saturation, to optimise signal/noise, although aiming for high resolution in a short scan - for instance in vivo - may sometimes force you to go below this level.

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Decreasing rotation step and increasing frame averaging will both strengthen the signal-to-noise ratio, at the cost of increasing scan duration. This is a trade-off that must be found for each bone scanning application.

Associated with the problem of noise in micro-CT imaging is the imaging artefact called “mottle”. As the name suggests, this implies a mottled or spotty appearance of images when looked at on a fine scale - for instance by “zooming in” on a micro-CT image on a PC screen to a high magnification. Mottle is a consequence of the random or “stochastic” distribution of x-ray photons striking the discreet pixel elements on the x-ray camera. It is a recognised feature of radiographic images19 and affects micro-CT due to the low flux from the micro- focus x-ray source. Mottle is illustrated in figure 7 which shows a magnified small part of the reconstructed crossection of a large bone sample - a human vertebral body. To quantify the degree of mottle, if you take the smallest image element discernible by the viewer and calculate the number of x-ray photons absorbed and detected within this element, the square root of this number (as a percent of the original number) represents the degree of mottle. Image noise in general including mottle is reduced by increasing the strength of the image signal. It can also be suppressed by image smoothing although this will cause some image detail to be lost.

Figure 7. Zooming in on a part of the reconstructed image of a scanned large bone (human vertebral body), mottle is evident, particularly in the low density background part of the image.

2.2.3. Ring artefacts

Some imaging artefacts are peculiar to CT, such as ring artefacts and beam hardening. Ring artefacts arise from the incremental rotation of the object- camera geometry, in combination with non-uniformity of the x-ray detection

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performance of individual camera pixels. Background subtraction, which is done to obtain a uniform image background (sometimes referred to as the “flat field” correction) should minimise ring artefacts. During image reconstruction ring artefacts can be suppressed by a Fourier analysis-based technique. Another “hardware” approach to ring artefact suppression involves linear movement of the x-ray camera or sample during acquisition so that a given sample voxel is imaged by several different camera pixels. This averages out differences in the output of a single pixel. Indeed some CT imaging systems employ continuous camera movement between each frame, which effectively reduces ring artefacts and also can increase the virtual size of the camera. The down-side to this however is some loss in image resolution due to limited camera movement precision, and significantly longer scan times. Ex-vivo micro-CT can make a limited number of random vertical sample movements during a scan, with a compensatory shift in the captured image by a few pixels. This has the same effect in averaging camera pixel output to suppress ring artefacts, with a small penalty in scan time and in image resolution. However such measures are more difficult to implement in in vivo scanners, since the large horizontally moving sample beds cannot be moved with the required precision, that is without unacceptable loss of image resolution. In vivo scanners with large cameras sometimes employ pixel binning (2x2, 4x4 for example) for faster acquisition at larger nominal resolution - this has a similar effect to camera movement in reducing ring features linked to aberrant camera pixels.

Figure 8. Soft ring artefacts caused can be efficiently removed by software Fourier based ring reduction - here the rings in the low density part of the left image are removed by software ring reduction.

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Ring artefacts can be categorised generally as “soft” or “hard” artefacts. Soft ring artefacts produce a series of concentric low contrast rings which can be more or less completely removed by Fourier Transform based image processing (figure 8). They arise from relatively low level variation in imaging performance between camera pixels. Hard ring artefacts can arise from dense objects in the uncorrected micro-CT image field caused for example by dense (e.g. metallic) dust particles somewhere in the beam path - such as in the sample chamber (if not cleaned regularly). These cause sharper ring artefacts which cannot be removed by conventional Fourier-based ring reduction - although other software approaches to creating a “dead pixel mask” can be effective. Furthermore, highly dense objects within the scanned sample itself, or or boundaries of dense material running parallel with concentric rings around the scan axis, can themselves cause local ring artefacts in the image. Thus the subject of ring artefacts is a complex one as there are multiple sources of such artefacts, and it is outside of the scope of this article to give a full treatment to all of these. In general the problem of ring artefacts becomes more severe with more x-ray dense samples; in scans of samples with low to intermediate x-ray density, which includes most bone samples (except very large ones), ring artefacts can usually be removed both by software ring reduction, and by keeping the scanner clean of heavy dust particles. Also, frequent flat field correction will reduce the number of ring artefacts, both soft and hard.

2.2.4. Beam hardening

Beam hardening is a fundamental problem of micro-CT. It results from the emission from the micro-focus x-ray source of a mixture of low and high energy x-rays. In an ideal world, CT imaging should be done with an x-ray source that is “monochromatic”, that is, it emits x-rays which have one energy only. But where the x-rays have mixed “polychromatic” energy, then as the x-rays pass through a solid sample, the low energy photons are removed from the beam preferentially so that the remaining beam has its energy spectrum shifted progressively higher. This process is called beam hardening since in x-ray physics parlance, “hard” means high-energy. Synchrotron CT can exploit monoenergetic x-rays, as can clinical CT to some extent by special techniques of x-ray generation. But these techniques are unavailable to micro-CT since the high intensity of the micro- focus source spot dictate the use of a source target - tungsten - which produces a broad range of x-ray energy (figure 9). This represents a compromise inherent to micro-CT imaging.

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Figure 9. The wide polychromatic spectrum of x-ray energy from the tungsten target of micro-CT scanners’ micro-focus sources is a source of the beam hardening artefact. It can be minimised by mathematical adjustment of the reconstruction algorithm, and by correct choice of filter and energy for the x-ray source.

Beam hardening is bad for CT imaging. The mathematical reconstruction of a CT image is usually based on the Feldkamp algorithm20 which consists of back- projection and convolution. Back-projection derives object position from the shadow position from multiple angles, while convolution removes a blurring halo from around the periphery of objects which back-projection alone would leave, by a selective image subtraction method. The Feldkamp algorithm is based on an assumption of equal attenuation along an x-ray path through the sample (in an exponential manner, as mentioned above), so that the signal from any voxel is independent of the ray path21. Beam hardening destroys this assumption, since the x-ray beam in its first interactions in its traversal through the sample will undergo greater attenuation than in its later interactions further through the sample, due to the increasing average x-ray photon energy.

Beam hardening can be minimised by software and hardware means. A variable correction can be applied to the reconstruction algorithm to compensate for beam hardening effects. On the hardware side, the spread of x-ray energy can be restricted from the low end by applying a metal foil in front of the source to remove the lowest energy part of the spectrum, and the high end restricted by lowering the applied voltage. However sometimes the lowest energy x-rays are required for imaging low density objects (such as soft biological tissue), and likewise the high end x-rays are needed to image dense objects. Beam hardening can to a varying extent be minimised but never entirely removed from micro-CT images. The problem of beam hardening is dealt with further in section 4 concerning density analysis, and illustrated in figure 13.

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2.2.5. Size of the scanned object

For objects composed of a material with a given general x-ray absorption - such as trabecular bone samples, as the size of the whole object increases, then the image of individual components within the object such as individual trabeculae will deteriorate - if scan duration and other parameters remain similar. To clarify this, say you take a large volume of more or less uniform trabecular bone from a human or sheep, and cut it into a large and a small sample, and scan both separately. With the same scan time and the same magnification (pixel size), the individual trabeculae will look clearer in the smaller sample scan. As a result, you generally have to scan longer to obtain a given image quality of individual trabeculae for larger samples of trabecular bone. There are several reasons for this. One is simply that for smaller samples you can employ a higher magnification and resolution (smaller pixel). A second reason relates to a parameter that one could call the “fractional absorption of a structural element” (FASE). In the above example of a small and large piece cut from the same trabecular bone sample, an x-ray beam travelling through the smaller sample will cross fewer trabeculae than going through the large sample. So the x-ray absorption of each trabecular structure is a larger percent of the total absorption in the projection x-ray image. This has the effect of increasing the signal-to-noise ratio of the image of that individual trabecular structure in the scan of the small piece compared to the large piece. (For the same reason, any container or other material surrounding a scanned sample - such as water - will decrease the signal-to-noise quality of the reconstructed image, if the surrounding absorption is significant compared to the sample absorption of x-rays.)

A third reason concerns the rotation of the scanned object. During scanning the object rotated around an axis that should be near the object’s centre. Close to this axis, rotation movements between image acquisitions (typically of less than one degree) cause a small movement in terms of absolute distance, while if you move radially further away from the rotation axis, parts of the object move a larger distance with each rotation. This effectively reduced the spatial resolution of the scanned images of parts of the object further away from the rotation axis. The practical result of this is that, for objects of larger size, a smaller rotation step is required to achieve a given spatial image resolution in a micro-CT scan of an object. Figure 10 illustrates this fact by superimposing a radial grid on a reconstructed crossection of a scanned bone.

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Figure 10. A radial grid superimposed on a reconstructed crossection of a scanned human vertebral body, illustrates the larger movements associated with scanning rotation at greater distances from the object centre.

2.3. Sample preparation and mounting for scanning

In general little or no preparation of bone or other samples is needed for micro- CT, reflecting the non-destructive character of the technique. However a few guidelines should be observed. (Here the mounting of samples for ex vivo scanning will be discussed - the setting up of animals for micro-CT imaging in vivo will be covered in a later section.) When harvesting bones from freshly sacrificed experimental animals (such as rats or mice) the bones should be cleaned as much as possible of soft tissue - although total cleanness is not necessary. One effective expedient for this is to take laboratory paper tissue, of the coarse, rough variety rather than fine and smooth, and use it to grab and pull muscle off the bone. When preparing the tibia (especially from a mouse) a sign of proficiency in this task is to be able to clean the bone of muscle without breaking the fibula (a formerly neglegted bone but becoming more important as a site for submicron nano-CT imaging, of osteocytes for example). Once dissected out and cleaned, the bones can be fixed in 10% buffered formalin for one or more days, followed by long term storage in 70% ethanol.

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Storage in alcohol for dry scanning is acceptable if the intention for micro-CT bone analysis is to study the trabecular or cortical morphology. However if a calibrated measurement of volumetric bone mineral density (vBMD) is required from micro-CT analysis, then dry scanning (and alcohol storage) are not adequate. The bones should be harvested and cleaned in the same way, but then placed in physiological saline and stored frozen. It is sufficient to wrap the bones in saline-soaked gauze.

When mounting a sample for micro-CT imaging, one obvious golden rule is that, apart from the controlled rotation of the whole object during scanning, there must be no movement of the object or any part of it. An example of movement that can compromise a scan is drying-related shrinkage. This will affect any normally hydrated soft biological tissue, so when scanning any sample in which soft tissue is exposed on the outside, it is necessary to isolate the soft tissue surfaces from contact with the air. A good solution for this is to wrap such samples in thin plastic film, such as the “cling film” obtainable from a supermarket. Generally however if a bone sample is more or less clean of soft tissue, and is scanned dry (taken from alcohol storage) then shrinkage will not be an issue. In this case the bone should be dried in air for 15-20 minutes on paper tissue prior to scanning. However if bones are to be scanned for vBMD, and are stored in saline and frozen as mentioned above, then they should not be dried prior to scanning (except superficially) but wrapped in plastic film in the same way as a soft tissue sample.

To prevent other movement during scanning, the sample needs to be held securely. Micro-CT users have improvised many good solutions to this employing for example plasticine (although this leaves a small residue on the sample), “parafilm” and plastic or expanded polystyrene tubes. As will be seen later (figure 14) any material surrounding the scanned object which absorbs x-rays significantly will decrease the quality of the object’s reconstructed CT image. In practice this is more of an issue for mouse than for rat bones due to the small size of the former. It means that if the sample is mounted in a tube, this should be of low density plastic such as polypropylene and not too thick-walled. Mounting in a fluid medium (water or alcohol) or embedding in resin such as methacrylate are possible, since these media absorb x-rays much less than calcified bone, but the imaging results will not be as good as when the bone is scanned dry and un- embedded. Another consequence of resin embedding is to limit the attainable image resolution, by increasing the size of the scanned object, as well as complicating the task of aligning the bone axially for scanning.

Another important issue, as just mentioned, is symmetrical rotation during scanning and alignment of the object around the (vertical) axis of rotation. There should be as little “wobble” from side to side as possible during the scanning rotation. Vertical axial alignment minimises the effective diameter of the sample and therefore increases the magnification available for that object - that is, allowing a smaller pixel size. The axial alignment of some frequently analysed bone samples, such as the rat femur, is complicated by the curved, “banana” shape of the bone. One solution to this (surprisingly effective in practice) is to mount the rat femur on an asymmetric (eccentric) mounting disc. With the bone correctly oriented on this disc, the result is that while the ends of the bone rotate

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away from the central axis, the section of interest that is imaged, the growth plate and metaphysis, rotate centrally and symmetrically in the camera field.

It is possible to mount several bone samples serially in a plastic tube, depending on the height of the scannable part of the sample chamber of the micro-CT instrument, and scan them sequentially in an automated, batch scanning operation. This is useful where large numbers of bones need to be analysed.

2.4. Three stages of micro-CT bone analysis: scan, reconstruct, analyse

The micro-CT analysis process can be described as having three stages: first the physical scan, acquiring the projection images from multiple angles; second the reconstruction of a stack of crossectional images that represent the 3D image “dataset”, by the Feldkamp algorithm for example; and third, obtaining quantitative data from a selected part of the image dataset. We can now turn to the latter, “tertiary” stage of data analysis that is probably of most interest to the bone scientist.

Having scanned your bone sample (in vivo or ex vivo) and reconstructed it to obtain the crossectional image dataset, the question is then what to do with the large volume of data thus provided. In practice there is not a narrow, standard analysis that must be performed - in fact there are nearly as many way to analyse micro-CT bone image data as there are research groups carrying out bone micro-CT, reflecting the versatility of the methodology. However, two of the most important categories of analysis are morphometric and density analysis, which will be briefly discussed.

3. Morphometric analysis of bone by micro-CT Morphometric analysis of bone includes analysis of trabecular and cortical bone for the same architectural parameters that are derived from histomorphometry (as discussed above). The crucial difference is that the micro-CT analysis is based directly on 3D image data. Important parameters in common with histomorphometry include relative bone volume (BV/TV), trabecular thickness (Tb.Th) or cortical thickness (Ct.Th), trabecular separation and number (Tb.Sp, Tb.N), surface to volume ratio (BS/BV) and trabecular pattern factor22 (TB.Pf) - the latter can be calculated in 2D or 3D. Other parameters have been developed from micro-CT analysis of bone which are fundamentally 3D quantities and have no equivalent in 2D-based histomorphometry. These include the structure model index (SMI) which indicates the relative prevalence of plates and rods in trabecular bone, indices of trabecular connectivity including the 3D calculation of trabecular pattern factor (TB.Pf) and the Euler number (Eu.N)23, and the degree of anisotropy (DA)24, 25.

Even parameters such as thickness which are also measured by histomorphometry, are measured by micro-CT analysis in a completely different way independent of any model assumption, such as “plate” or “rod”. Trabecular

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thickness for example can be measured by a method called “local thickness” developed by Hildebrand et al. 10 and requires a 3D model of the bone. Local thickness for a point in the bone part of the model is defined by these authors as the diameter of a sphere which fulfils two conditions: (a) the sphere encloses the point (but the point is not necessarily the centre of the sphere), and (b) the sphere is entirely bounded within the solid surfaces. This analysis can provide a distribution of thicknesses as well as a single mean value of Tb.Th from a trabecular bone site. A further strength of the local thickness method is its independence of orientation of the studied object.

There are two steps that need to be carried out with a reconstructed “dataset” - the stack of crossectional images - in order to perform morphometric analysis. These are (1) binarisation, otherwise known as segmentation, and (2) selection of the volume of interest. The first, binarisation, is illustrated in figure 11. The grey level images produced by reconstruction consist of voxels which have 256 grey levels (if the images are 8-bit). The binarised image is 1-bit with for example black representing solid, or bone, and white representing space. The simplest and quickest method - and most widely used - is “global thresholding”. This means setting a single threshold grey level density value, so that every voxel with an equal or higher value is represented as solid colour, and lower values represented as space. Global thresholding can be refined by some image analysis techniques such as “despeckle” to remove single dot objects that are most likely to arise from noise. Global thresholding is generally adequate for obtaining morphometric data where the bone images are of sufficient quality, as they should be from most ex-vivo imaging applications.

However in the special case of in vivo imaging of bone, images obtained will have a lower signal-to-noise ratio. This is for several technical reasons, including the additional x-ray absorption of soft tissue around the bone, the short duration of the scan necessitated by in vivo experimental and animal welfare considerations, and the less flexible geometry of in vivo micro-CT scanners. Waarsing et al. 26

describe the software segmentation technique of “local thresholding” which improves the separation of bone from non-bone in noisy, in vivo images. This technique does not impose a global threshold value but calculates threshold values locally at different locations in the bone image dataset on the basis of bone surface gradients of grey level that are related to the thickness of the solid structure - such as a trabecular plate. The advantages of this local thresholding are noise suppression, better preservation of trabecular connections, and improved ability to segment both high and low density bone regions in the same micro-CT image.

3.1. Calibration of micro-CT bone morphometry (trabecular thickness)

A few words should be mentioned about calibration of bone morphometric measurements such as thickness. In the majority of published papers on bone analysis by micro-CT to date the question of calibration has been ignored. However, as Ito et al. 27 first pointed out, the structure thickness that you obtain from a micro-CT measurement is very sensitive to the selected threshold value, especially when using simple global thresholding. It is therefore necessary to calibrate thickness measurement if one wishes to confirm that measured values are somewhere close to actual bone structure thicknesses.

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Calibration of bone thickness can be made using aluminium foils, and is necessary whether you are using global or local thresholding. Aluminium provides a suitable material for calibration of micro-CT measurement of bone structure thickness. Although Al is more physically dense than bone (2.6 g cm-3

compared to 1.3-1.8 g cm-3) the higher atomic number of Ca (20) compared to Al (13) compensates for this owing to the acute dependency of x-ray absorption on atomic number, as discussed above. The x-ray opacity of aluminium is as a result very close to that of cortical bone. Aluminium is also materially uniform on a micron scale, which is important for precise calibration of thicknesses in microCT systems with image resolution in the order of a few microns.

Some thickness measurements of aluminium foils with thicknesses of 20 and a 250 µm (Goodfellows Ltd., England) are shown in table 1. (The scanner used was a Skyscan 1072). Structure thickness was measured by the sphere fitting local thickness method10. Note that the foils thickness as quoted by the manufacturer is known only with a precision of ± 10%. This limits the accuracy of the calibration.

The foils were scanned at a range of nominal resolution (pixel size) settings. In table 1 the measured thicknesses of the foils is given as a function of a range of micro-CT magnifications and pixel sizes from 3.9 - 11.7 µm. As the pixel size increases from 3.9 µm for the 20 µm foil measurement, the measured thickness remains approximately correct until a pixel size of 9 µm, at which point thickness becomes exaggerated by about a factor of two. A conclusion from this is that accurate thickness measurement of a structure such as aluminium foil or bone is possible only where the structure thickness corresponds to three or more times the scan pixel size.

Aluminium wires can also be used to calibrate measurement of rod-like trabecular bone structures. Such rods or wires are more demanding to measure for thickness than foils (due to shorter x-ray paths through the wire), and accurate measurement requires a minimum thickness of at least 5-6 pixels.

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Table 1. Measurements of thickness of aluminium foils with 20 and 250 µm thickness, to calibrate trabecular bone thickness measurement. Measurements of the 20 µm foil are approximately accurate at 3.9 µm up to 6.8 µm pixel size but not at pixel sizes above 9 micron. This suggests a minimum structure thickness of about three times the scan pixel size is needed for accurate measurement by micro-CT.

Micro-CT pixel Foil Magnification Foil thickness Foil thickness size (“nominal thickness measured, mm in pixels resolution”) mm

11.7 0.020 x 23 0.04115 1.71

9.1 0.020 x 30 0.03925 2.20

6.8 0.020 x 40 0.02029 2.92

5.5 0.020 x 50 0.02124 3.64

3.9 0.020 x 70 0.01922 5.13

11.3 0.250 x 24 0.2362 22.12

4.6 0.250 x 60 0.2481 54.35

3.9 0.250 x 70 0.2379 64.10

Selection of the region, or volume of interest (VOI) is the starting point of morphometric analysis. (Binarisation can be either before or after VOI selection). An important and common example is the selection of a volume of trabecular bone at a standard trabecular site in a rodent bone, such as the distal femur metaphysis (the trabecular bone below the growth plate at the knee end of the femur) or the proximal tibia. For morphometric measurements it is useful to repeat the emphasis made by Parfitt et al. 13 on the central importance of the “referent” for histomorphometric measurements which calculate a relative value. For example, for relative trabecular bone volume (BV/TV) the referent is TV, the tissue or total volume. The referent is the area or volume within which the measurement is performed, and trabecular BV/TV has no meaning unless measured within a referenced volume of medullary bone containing trabecular bone and marrow only, which should not extend into cortical bone or out of the bone tissue. Again the new aspect of micro-CT analysis compared to histomorphometry is that the VOI is in three dimensions.

In practice this means that in your micro-CT analysis software you need to scroll through the multiple crossection levels at the trabecular region you wish to study, and draw at several separated levels the region you wish to analyse (refer to figure 12). This allows a complex polygonal 3D volume of interest to be selected. Such software packages generally allow the crossectional images to be resampled into longitudinal (“sagittal”) images of the type more familiar to histomorphometrists, giving the option to delineate the VOI from longitudinal images, where the ability to see the growth plate and metaphyseal trabecular

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bone in the same image may be an advantage. This is really a question of preference of an individual analyst. Figure 11. As a pre-requisite for morphometric analysis of micro-CT images, the grey- level image data must be transformed to binary images distinguishing bone from non- bone, a process called binarisation or segmentation. High resolution images such as ex vivo bone images can adequately be binarised by global thresholding (a and b: mouse femur ex vivo). More noisy rapidly scanned in vivo images (c and d) may require more advanced segmentation techniques such as the “local thresholding” developed by Waarsing et al. 22.

Automated software routines can be used for separating trabecular from cortical bone. However at some bone sites, such as close to the growth plate or in vertebrae, the cortical bone is thin, perforated and poorly defined, so that automated methods are limited in application. Manual VOI delineation has the advantage of providing quality control for the scanned datasets, and letting the analyst actually look at the imaged bone structures and gain a visual impression of the experimental outcome (although the analyst should of course be “blind” as to the sample’s identity).

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Figure 12. Selecting the region - or volume - of interest (VOI) is the starting point of morphometric analysis. A volume referent is required for many morphometric parameters 13. Here the selection of a metaphyseal trabecular VOI is illustrated from an ex vivo micro-CT scan of a mouse distal femur. (a) a rendered model of the whole scanned distal femur, (b) software selection of the trabecular VOI on several crossections, and (c) a rendered model of the selected volume of the distal femur medulla.

4. Density analysis of bone by micro-CT

Scientists looking at micro-CT images of materials such as bone, are often attracted by the possibility of obtaining detailed density data in 3D, by interpreting the grey level of each voxel as the density at that point. However, again, caution must be exercised in interpreting micro-CT density information, since the density information interpreted by this voxel-by-voxel approach is subject to errors and artefacts, and is thus frequently simply inaccurate. Three CT imaging artefacts (at least) influence the reported density at any voxel. These are the partial volume effect, beam hardening (which has already been discussed) and the thickness of surrounding material. The first two phenomena are illustrated in figure 13.

The partial volume effect means that objects below a certain thickness, close to the minimum resolvable thickness, have their density under-reported. This is shown in figure 13 for aluminium foils. The reason for it is that at interfaces between low and high density parts of a CT image (for instance at an object surface), there is a gradient of voxel density with a finite slope - as seen in the profile lines in figure 13. For a thin object the density gradients from opposite sides of the object meet before they attain the true material density value.

What this means for the bone micro-CT analyst is that the volumetric, material density of trabecular bone cannot accurately be measured from micro-CT images, unless a complex correction is made for the decrease in apparent density with changing thickness - using for example aluminium foils, and involving the measured thickness distribution of the trabecular bone. The apparent density in micro-CT images within trabecular bone structures generally reflects the trabecular thickness at that point more than the true material density. However it is more practicable to measure the density of a selected volume of bone medulla containing both trabecular bone and marrow. Such a measurement

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should be termed something like “regional” mean bone mineral density, to distinguish it from actual volumetric calcified tissue density. The latter could by contrast be measured accurately for cortical bone (or thicker trabecular plates scanned at high enough resolution) where the partial volume effect was small enough for a true, flat-bottomed density profile to be obtained - as in the two thicker foils in figure 13 (a).

Beam hardening as shown in figure 13 (b) artificially increases the density of surface layers of a scanned object. It is easier to see the effect of beam hardening where sufficiently large areas of solid material - such as cortical bone - are imaged. In practice the software correction for beam hardening is optimised by trial and error, and the degree of correction is selected which in the analyst’s view eliminates the artefact. This procedure can be helped by use of image viewing and analysis software which shows density profiles - such as shown in figure 13. For any scans of bones intended to be used for measurement of density it is very important to optimise the correction for beam hardening. Furthermore, as discussed above, beam hardening can be minimised by careful selection of x- ray imaging filter and applied tube voltage.

(a) (b)

Figure 13. Density measurement in 3D by micro-CT is complicated by several artefacts. The partial volume effect (a) is reflected in the density profile through four aluminium foils, with thickness (from left to right) of 250, 20, 50 and 125 µm. The profile of x-ray absorption is shown beneath - lower values mean more absorption and higher density. While the actual density of Al the foils is the same, the two thinner foils show lower density in profile. Beam hardening is shown (b) in two reconstructed crossections of the same image slice from a rat femur. The upper image is not corrected for beam hardening while the lower image is corrected. Beam hardening is evident as an artificially increased density in a thin surface layer at the periphery of the scanned object.

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Thirdly, the reported density of a scanned object is significantly affected by the thickness of other material surrounding that object. The surrounding material effectively acts as additional x-ray filtration. As a result, the same object or material scanned with little or no surrounding material will have a CT image density higher than when surrounded by a thick layer of a material. This is illustrated in figure 14, which shows progressive crossection through a mouse femur imaged in vivo (Skyscan 1076), from the proximal (hip) end toward the distal (knee). With this progression, the thickness of surrounding soft tissue steadily decreases, and the bone’s apparent density correspondingly increases. (This is an artificial, not a true increase in bone density.) Figure 14 also shows that the quality of the image, in terms of signal to noise ratio, improves with a reduction in the thickness of surrounding soft tissue. For the same reason, the best quality micro-CT images of bone are obtained ex vivo, scanning a cleaned bone with no surrounding layer.

However, it should be stated that with careful experimental procedure, all three of these sources of error can be minimised sufficiently to allow bone mineral density to be measured from micro-CT images. An important part of this method is calibration using phantom materials with known bone density, defined for example as the density in g cm-3 of calcium hydroxyapatite. Both the phantom bone equivalent material, and any surrounding soft tissue equivalent material (such as Perspex or water) should have as closely as possible the same thicknesses as the bone and surrounding soft tissue (if any) being scanned.

5. Ex-vivo application of bone analysis by micro-CT

Bone by its physical and chemical nature is very well suited for x-ray micro-CT analysis, as discussed previously. No attempt will be made here to exhaustively review all bone micro-CT applications - there are now too many for this to be practicable. To illustrate the steadily increasing use of micro-CT in bone research, an internet biomedical database search by PubMed was made for the words “bone” and “micro-CT”, restricted to a single year for the years 2001- 2006. The number of hits for these years respectively was 12, 19, 26, 45, 56 and 82. This search no doubt missed some published micro-CT bone studies, but the trend is clear. Bones of rats and mice, at standard sites for histomorphometry such as the distal femur, proximal tibia and lumbar vertebra, are very suitably sized for analysis by commercially available micro-CT scanners, whose width of field of view is typically 2-4 cm. Furthermore, biopsies of bone from human patients, taken from the posterior iliac crest, are generally about 8mm in diameter, also ideally sized for micro-CT. Some examples of osteoporosis-related research involving ex vivo scanning of rodent bones, and of human biopsied bone, will be described here, emphasising the unique advantages provided by non-destructive micro-CT 3D analysis. (There is a bias toward papers in which the author was involved.) In the subsequent section the newer field of in vivo scanning of bone by micro-CT will be looked at.

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Figure 14. The thickness of surrounding soft tissue affects the reported density of bone in an in vivo scan. Here progressive crossections are shown through the in vivo imaged mouse femur, from proximal (hip) end toward the distal (knee). As the thickness of surrounding soft tissue decreases, reported density of the bone increases (this is an artefact, not a real change in bone mineral density). Also, the image signal to noise ratio improves.

The efficacy of two bone anabolic agents, the parathyroid hormone fragment hPTH(1-34) and fibroblast growth factor (bFGF) , was studied by Lane et al. 28 in female ovariectomised rats. Micro-CT morphometry was employed to assess the murine trabecular architecture. This study demonstrated the ability of micro-CT to quantify three dimensional parameters of bone in a way not possible by 2D histomorphometry - specifically trabecular connectivity. This 3D connectivity measurement turned out to be the central story of the experiment - both anabolic agents increased trabecular volume, but only bFGF enhanced the number of inter-trabecular connections, in a manner suggesting useful potential for therapeutically reversing the loss of trabecular connectivity associated with severe osteoporosis in old age. By contrast hPTH(1-34) achieved an enhancement of trabecular volume through an increased trabecular thickness

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but without increased connectivity. The study by Lane et al. illustrates well the advantages and new analytical power of micro-CT morphometry compared to 2D histomorphometry.

Figure 15. Micro-CT analysis of the male mouse distal femur. (a) and (b) show trabecular bone from sham operated and orchidectomised mice. Treatment with estrogen (c) increased both trabecular volume and thickness, while testosterone treatment (d) enhanced trabecular volume by increasing trabecular number density without any effect on trabecular thickness.

Moverare et al. 29, also in 2003, compared the architectural response of trabecular bone from male orchidectomised mice to estrogen and testosterone. Micro-CT morphometry was applied to the mouse’s distal femur trabecular bone. This study was done in two parts, one with wild type and the other with double estrogen receptor knockout (DERKO, alpha and beta receptors absent). In the wild type mice, both estrogen and testosterone protected the mice from ORX- induced bone loss. However the mechanism of trabecular volume enhancement by the two sex steroids was different: estrogen treatment increased both trabecular number and thickness, while testosterone increased trabecular number only (figure 15). In the DERKO male animals, estrogen had no effect on trabecular volume - confirming successful removal of estrogen receptor effect - while testosterone fully protected trabecular volume. This showed that testosterone was acting through an androgen, not estrogen receptor (plus aromatase) route. Interestingly in the four DERKO groups, there were no significant differences in trabecular thickness, unlike among the WT groups. This finding indicates an association between the estrogen receptor and changes to trabecular thickness, raising some mechanistic questions. A feature of this study was that the mouse bones, following micro-CT analysis, went on to be analysed further by biochemical techniques.

In another study from the same group, on sex steroid bone biology, Tivesten et al. 30 used female rats which were ovariectomised or sham operated. The distal femoral metaphyseal trabecular bone was again studied by ex vivo micro-CT morphometry. OVX rats received either vehicle, estrogen, testosterone or both sex steroids together. Both steroids separately did significantly enhance trabecular volume of OVX rats - estrogen being substantially more effective - but estrogen and testosterone together were significantly more efficacious than estrogen alone. The interesting aspect was that this additive efficacy was reflected in morphometric parameters - trabecular thickness and number. Unpublished additional data from this study show that the effects of estrogen

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and testosterone were also additive in other measured 3D architectural parameters, namely the structure model index, trabecular pattern factor and Euler connectivity. This study has implications for the possible clinical use of hormone replacement therapy employing both estrogen and testosterone based compounds.

In addition to “involutional” osteoporosis related to sex hormone deficiency, micro-CT has proved valuable in the study of glucocorticoid induced osteoporosis. Salmon et al. 31 reported the development of a murine model for glucocorticoid-induced osteoporosis. This again employed the femoral distal metaphysis, analysed ex vivo by micro-CT. Glucocorticoid induced osteoporosis (GIOP) has proved difficult to reproduce in a mouse model. Mice frequently gain increased trabecular relative volume in response to glucocorticoid administration. Murine vertebral cortical bone was evaluated as an alternative to trabecular bone as a model of GIOP.

Mice were treated by injection with either vehicle or the steroid prednisolone for five weeks (see figure 16). Lumbar vertebral morphometry was assessed post- mortem by high resolution micro-CT with a pixel size of 5 µm. Trabecular bone relative volume (BV/TV) at the mouse lumbar vertebral body showed biphasic changes which inversely mirrored the overall vertebral volume. BV/TV significantly increased in the group receiving low and intermediate doses of prednisolone, while higher doses showed decreasing BV/TV. Bone thickness showed a more unidirectional response. There was a significant dose-dependent reduction in both trabecular and cortical thickness over the whole dose range, more pronounced and significant for cortical bone. The higher trabecular volumes at intermediate prednisolone doses were associated with more plate- like structure (a lower structure model index - see appendix). Unlike trabecular bone, cortical thickness and volume showed a dose-dependent decrease in response to steroid over the whole dose range. Cortical bone therefore appears to be superior to trabecular bone, in the mouse vertebra, as a model indicator of glucocorticoid-induced osteopenia in humans caused by the steroid-induced reduction in bone formation.

McLaughlin et al. 32 also used micro-CT to study effects of the glucocorticoid dexamethasone on murine bone. The above studies indicate the power of microCT bone morphometry to go further than just indicating gain or loss of bone volume, to elucidate subtle mechanisms of change in bone volume and architectural patterns of change.

Another skeletal site that can be used to assess anabolic effects on bone formation is the calvaria - one of the bones comprising the roof of the skull. The mouse calvaria has been established for some years as a site for assessment of anabolic action of small amounts of an agent by the local injection of the substance onto the mouse calvaria (e.g. 33). This can be useful when a substance under investigation is not readily available in large enough quantities to be used in a normal animal study assessing effects of systemic injection.

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(a) (b)

Figure 16. The changes in trabecular and cortical morphometry in a mouse model of glucocorticoid induced osteoporosis and smaller and more subtle than those associated with experimental gonadectomy to simulate osteoporosis. Micro-CT is a powerful analytical method to evaluate such changes. Prednisolone was shown to cause decreased thickness of vertebral trabecular and cortical bone in the mouse model 27 as shown above comparing a control vehicle-treated mouse (a) to a mouse after prednisolone administration (b).

Maeda et al. 34 employed the murine calvarian injection model to assess the bone anabolic effect of the inhibitory helix-loop-helix transcription factors Id1 and Id3. The two factors were found to be effective at promoting new bone growth. The mouse calvaria were imaged by micro-CT post-mortem (although this can equally be done in vivo). The advantage here of micro-CT is that, since the calvaria are imaged and analysed in a fully 3D environment, virtual sections could be taken which were precisely at right-angles to the calvarial plane, improving the precision of the growth rate measurement.

In addition to pharmacologic effects on bone, there has been substantial research into nutritional effects on bone architecture and osteoporosis. The changes to bone architecture from nutritional factors in an experimental context can be smaller and more subtle than from pharmacological agents. This is another application in which micro-CT bone morphometry is advantageous due to its power to detect small changes in bone morphometry. Analysing an entire metaphyseal trabecular region for example, in 3D by micro-CT, provides a more powerful and precise structural analysis than histological measurements from one or two 2d sections. Medeiros et al. 35 employed micro-CT to study the effects of iron deficiency on lumbar vertebral trabecular architecture in female weanling rats (using pair-feeding to normalise dietary intake). Iron deficiency was shown to cause wide-ranging architectural deterioration: trabecular volume was reduced, due to a reduction in both trabecular number and thickness. Structure model index also increased indicating a pathological change from plate-like toward rod-like structures. These effects of iron deficiency were lesser than equivalent effects of calcium restriction: again micro-CT was effective in detecting rather subtle but important architectural changes.

In addition to osteoporosis, micro-CT analysis has been shown to be a powerful tool to precisely quantify bone changes from other pathologies, such as bone

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tumour 36, myeloma 37 and collagen induced arthritis 38. It is also becoming an increasingly important methodology in orthopaedic research, for example for assessing synthetic biomaterials and bone implants. The large field of orthopaedic research, including the application of finite element analysis for virtual strength testing of micro-CT images, is outside the scope of this article. However one comment should be made regarding the imaging of metallic implants in bone. Metals can cause special problems for micro-CT imaging due to their high absorption of x-rays (refer to the earlier section on photoelectric absorption). This is related to their atomic number. Essentially what this means is that metals such as titanium and aluminium with relatively low atomic number can be successfully imaged in contact with bone by micro-CT, but any object composed of steel will cause such strong absorption of the x-rays (with the energies employed in micro-CT) that gross imaging artefacts will be formed around the steel objects in reconstructed images. So avoid steel if possible when conducting micro-CT imaging of orthopaedic implants.

As mentioned earlier, biopsied samples of human iliac crest trabecular bone are also highly suitable samples for micro-CT analysis. Won et al. 39 took biopsy samples from premenopausal women, which were analysed by micro-CT and bone histomorphometry (taking advantage of the non-destructive nature of micro-CT imaging). Measurements of morphometric parameters were quite well correlated between the two methods, providing validation of micro-CT clinical biopsy assessment, with respect to histomorphometry. An earlier validation comparing micro-CT morphometry to histomorphometry was published by Ito et al. 23. Validation studies have also been performed comparing the results of bone histomorphometry with synchrotron micro-tomography 40 and conventional clinical CT 41.

Finally, micro-CT has been employed in the study of bone biopsy samples from a teenager suffering from a severe genetic disorder of bone, called Juvenile Pagets Disease (JPD - or alternatively, Idiopathic Hyperphosphatasia). This condition is caused by mutations inactivating the gene for osteoprotegerin (OPG), and the bone pathology of an 11 year old girl with severe JPD has been described recently by Cundy42. The condition is debilitating and is characterised by severe osteoporosis associated with a greatly accelerated bone turnover rate. One unusual architectural feature of trabecular bone studied from iliac biopsies of the girl with JPD, and also seen in previous JPD cases, is a striking parallel arrangement of trabecular plates, in place of the normal randomly oriented pattern (see figure 17). As well as being unusual, this architecture is very poor biomechanically. A possible explanation for this phenomenon was proposed by Salmon43 involving the interpretation of formation-resorption coupling mediated by RANKL as dynamic “feedback”. To summarise, a consequence of the remodelling feedback (caused by OPG deficiency) is the suppression of the spontaneous nonlinear pattern formation that establishes normal chaotic trabecular architecture, and the imposition of regular parallel plate architecture. Physical experimental systems showing spontaneous pattern formation can also demonstrate suppression of complex pattern by feedback.

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Figure 17. Micro-CT generated 3d models44 of iliac crest trabecular biopsies from a JPD patient before (a) and after (b) bisphosphonate treatment, and a control normal individual (c).

Until the study by Cundy (2004) the abnormal JPD trabecular bone phenotype had only been studied by histology. Now however the biopsied JPD bone has been analysed by micro-CT (in as yet unpublished work), giving more information on the trabecular structure, before and after bisphosphonate therapy (figure 17). While histological images did not show any effect of bisphosphonate on the parallel trabecular structure, micro-CT analysis of a volume of iliac trabecular bone, pre- and post-bisphosphonate treatment, showed a significant reduction in trabecular anisotropy in the latter. So bisphosphonate is at least partially efficacious in ameliorating the abnormal parallel trabecular architecture in the JPD patients.

6. In vivo applications of bone analysis by micro-CT

A recent development in micro-CT technology (since 2002) is the availability of in vivo micro-CT scanners, which can image bones of a living, anaesthetised rodent (Sasov et al. 2002). These scanners have a different scanning geometry to ex-vivo scanners, in that the object remains stationary during a scan, while the x- ray micro-focus source and the camera are mounted on a gantry which rotates around the object. This is the same architecture as a medical CT scanner, but in vivo micro-CT scanners have maximum nominal resolution (minimum image pixel size) down to below 10 microns. Figure 18 illustrates the mounting of animals (in this case mice) for in vivo micro-CT analysis of several bone sites, and the results obtainable in vivo in terms of crossectional and 3D images of bone.

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Figure 18. Crossectional and 3D imaging detail obtainable from in vivo scans of the hindlimb, skull, forelimb and tail caudal vertebra, with scan durations from 10-25 minutes (Skyscan 1076). Use of custom-made polystyrene tubes, and plasticine to secure limbs or the tail, allow a mouse to be mounted for in vivo micro-CT scanning of a range of bone sites.

Several bone sites most commonly used in ex vivo preclinical bone analysis can be assessed in vivo. These include - as shown in figure 18 - the knee sites (proximal tibia, distal femur), the skull, the forearm and the caudal tail vertebrae. The knee sites are perhaps the most important sites in analysis of rodent bone, and are well suited to in vivo micro-CT imaging due to the relatively small thickness of overlying soft tissue around the knee. As a result, a scan of about 20 minutes duration can provide image resolution and quality sufficient to adequately image mouse or rat trabecular bone architecture (Waarsing et al. 2004b). The skull can also be imaged with a good image quality, allowing for

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example in vivo implementation of the murine calvarian local injection models (e.g. Izbicka et al. 1997).

The forearm can be scanned in a similar manner to the knee. This allows the ulna to be imaged in vivo, which allows in vivo implementation of ulna mechanical loading models which are currently assessed ex vivo 45, 46 Warden et al. 2004). Finally, the caudal vertebrae in the mouse tail provides an advantageous site for in vivo scanning of vertebral trabecular bone, due to the minimal surrounding soft tissue. The caudal vertebra is much more suitable for in vivo imaging than the lumbar vertebra - the most commonly studied vertebral bone ex vivo - since the latter is surrounded by thick abdominal tissues which seriously degrade the signal-to-noise quality of scanned images.

The animal should lie comfortably during a scan. The use of expanded polystyrene tubes as illustrated in figure 18 provides thermal insulation for a mouse, as well as minimising surrounding x-ray absorption. Temperature control - specifically preventing a scanned animal becoming cold during a scan - should also be effected by control of the airflow temperature as part of the physiological monitoring system of the in vivo scanner. Rats are too heavy to support in expanded polystyrene tubes, and carbon fibre beds about a millimetre thick provide good rigidity with low x-ray absorbance. Generally in vivo scanners provide video camera surveillance of your animal during a scan. Also available are a mask and air tube for breathing monitoring and ECG electrodes and cables to measure heart beat. However for bone scanning it is usually sufficient only to monitor the animal by live video image. The issue of synchronisation of image acquisition with the breathing cycle is most relevant to the imaging of lungs rather than bone, unless one is imaging the ribs or thoracic vertebrae. In practice the irregular nature of the breathing of a rodent under anaesthesia (especially injection anaesthesia) combined with the significant x-ray camera exposure time compared to the breathing cycle duration, mean that significant image improvement by synchronisation, from scans of the upper thorax, requires active control of the animals breathing cycle using a ventilator, a complex and invasive procedure. Some suppression of breathing movements in the upper thorax can be achieved by the much more low-tech means of wrapping paper-based adhesive tape around the upper chest (figure 18, second row), applying light pressure. This has the effect of driving the breathing movements in large part to the abdomen, away from the upper chest.

Owing to the recent introduction of in vivo micro-CT scanners, little is yet published in the open literature on bone micro-CT analysis in vivo. The first such study to be published was by Waarsing et al. 2004b, who impressively demonstrated the entirely new capabilities of in vivo micro-CT scanning, such as to track trabecular bone changes following ovariectomy in the female rat, at the level of individual trabecular structures (see figure 19). The rats proximal tibias were scanned at the time of ovariectomy (OVX), and then at 4, 14, 34 and 54 weeks after surgery, using in vivo micro-CT at a nominal resolution (pixel size) of 18 µm. The software technique of image registration - 3D matching of time- separated images of the same object - was applied, allowing conclusions to be drawn as to the tissue level mechanism of post-OVX trabecular bone loss. The authors found that initially following OVX, trabeculae are selectively lost in their entirety, dramatically reducing trabecular number. At a later stage (beyond 14

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weeks post-OVX) the few remaining trabecular structures undergo thickening, and some new trabecular structures grow to “bridge” gaps left by the early trabecular loss. Linear alignment of the remaining trabeculae develops above and below the growth plate, apparently to optimise mechanical strength along the bone’s long axis. The whole proximal tibia was shown to undergo a small change in shape in the weeks after OVX.

A comment should be made concerning radiation doses during in vivo micro-CT scanning. The radiation doses reported by Waarsing et al. (2004b), were 0.4 Gy for 20 minute scans. (In terms of equivalent dose, this corresponds to less than 5 mSv respectively using human tissue weighting factors from ICRP 6047. Radiobiological studies indicate that harmful effects to growing chondrocytes (which form new bone at the growth plate) begin to occur in the range 1-2 Gy48. However experimental studies on radiation induced bone growth retardation typically require substantially higher acute x-ray doses in order to obtain a measureable effect. Knee irradiation of 4 week old Sprague-Dawley rats with a single acute x-ray dose of 25 Gy reduced bone growth rate by 59% (Damron et al. 2001). Irradiation of the knees of 4-week old rats by a single 17.5 Gy acute x-ray dose transiently reduced bone growth rate at the proximal tibia by 18% (Damron et al. 2003). These doses are 62 and 44 times higher respectively than the 0.4 Gy associated with an in vivo knee scan in the Skyscan 1076 as described above. Extrapolation of a simple “dose-response” curve between these two studies, would suggest that doses to the rodent knee of 0.5 Gy or less should not provoke excessive concern about significant bone growth retardation.

Experience of in vivo scanning to date tends to confirm these findings - single scans of half an hour or less, or repeated scans with an interval of a month or more do not apparently affect bone growth. But scans separated by only a week or less cause reduction in bone growth (as assessed by comparing a scanned to a non-scanned leg) since the radiation dose becomes additive. More research clearly needs to be done to elucidate and quantify the effects of radiation doses received during in vivo scanning. It should also be borne in mind that general anaesthesia itself is a profound trauma to a rat or mouse, causing among other effects, significant growth retardation49, so in assessing the ethical or scientific value of in vivo scanning protocols, anaesthesia is an equally significant factor as radiation dose. Both should be minimised by adopting conservative protocols with the number of in vivo scans minimised, and their timing separated by several weeks. Incidentally, the final micro-CT imaging at the end of an in vivo study has no significance in terms of physiological effect of radiation dose or anaesthesia, since it coincides with termination of the study and sacrifice of the animal. It could be performed for example on the freshly killed animal.

The image registration methods employed by Waarsing et al. 26 require at present a high level of computational / mathematical expertise for laboratory realisation. However in vivo micro-CT can offer substantial advantages to bone research at a simpler level of implementation - simply scanning and analysing referenced regions of bone at more than one time point in an in vivo experiment (the usual growth-plate referenced trabecular sites for example). The ability to study a rat or mouse’s trabecular or cortical architecture at the start and the end of a study profoundly increases the statistical power and precision of in vivo bone studies by removing the assumption of homogeneity between animals at

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the start of the experiment. Each animal becomes its own control. The advantages of “before and after” in vivo bone scanning were demonstrated by Offley et al. 50 who investigated the role of unmyelinated peripheral nerves in maintaining trabecular bone volume and architecture in the rat. Rats were administered capsaicin, which destroys unmyelinated sensory neurons, or vehicle. These authors scanned the proximal tibia of rats by in vivo micro-CT. The tibias were scanned on the live animals before and after the dosing period. With a group size of 8, modest changes in volume and architecture were demonstrated with good statistical significance. It was confirmed that loss of peripheral sensory unmyelinated nerves caused loss of trabecular bone volume and deterioration of architecture.

7. Conclusion The future development of micro-CT in bone research will probably be characterised by a growth in the realisation of the special advantages of in vivo micro-CT, matching the already established expansion of ex vivo micro-CT in bone research laboratories. Micro-CT in vivo scanning is also being combined with other imaging modalities such as positron emission tomography (PET), single photon emission computed tomography (SPECT) and luminescence imaging. All three of these methods allow specific molecules such as administered drugs or tagged immunolabels to be imaged in 3D, for example within a living mouse or rat. Such combinations of imaging technologies are referred to under the term “molecular imaging”. The technology of micro-CT itself can be expected to advance significantly, both in terms of high resolution image production technology, and computer image and data analysis. Progress should be made in further minimising the inherent errors in the technique which were discussed above. Micro-CT should eventually gain acceptance as the “gold standard” methodology for obtaining 3D morphometric data on mineralised bone tissue.

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Figure 19. The pioneering use of in vivo micro-CT in bone research by Waarsing et al.51

Each micro-CT cross-section of the rat proximal tibia represents the same area of the bone at week 0 (A1, B1), week 4 (A2, B2) and week 14 (A3, B3). (A) An example of trabecular thinning in the sham-operated rat. (B) An example of thickening of the remaining trabeculae in the OVX rat.

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