Synchrotron Imaging Techniques for Bone and Cartilage Tissue Engineering: Potential,Current Trends,and Future Directions

Abstract

Biomedical imaging is crucial to the success of bone/cartilage tissue engineering (TE) by providing detailed three-dimensional information on tissue-engineered scaffolds and associated bone/cartilage growth during the healing process. Synchrotron radiation (SR)-based biomedical imaging is an emerging technique for this purpose that has been drawing considerable recent attention. Due to the unique properties of synchrotron light, SR biomedical imaging can provide information that conventional X-ray imaging is not able to capture. SR biomedical imaging techniques notably differ from conventional imaging in both physics and implementation, thus varying with regard to both capability and popularity for biomedical imaging applications. In the earlier decade, synchrotron-based imaging was used in bone/cartilage TE to characterize bone/cartilage scaffolds and tissues as well as the varying degrees of success in reconstruction. However, several key issues should be addressed through research before SR biomedical imaging can be advanced to a noninvasive method for application to live animals and eventually to human patients. This review briefly presents recent developments in this area, focusing on different synchrotron-based biomedical imaging techniques and their advantages and limitations, as well as reported applications to bone and cartilage TE. Key issues and challenges are also identified and discussed along with recommendations for future research.

Introduction

Three-dimensional (3D) biomedical imaging plays a major role in determining the efficiency of tissue-engineered regeneration. It enables an examination of microstructural, biochemical, and biological properties of tissue-engineered scaffolds and, once implanted, their interaction with host tissues.¹ X-ray absorption imaging (such as microcomputed tomography [μCT] and CT),² ultrasound,³ and magnetic resonance imaging (MRI)^1,4 are among the prevalent 3D imaging methods used for noninvasive characterization of cartilage and bone tissues, scaffold microarchitecture, and tissue growth. Although these techniques are currently in widespread use for clinical diagnosis, limitations preventing advanced tissue engineering (TE) applications include low spatial resolution,⁵ poor contrast of low Z elements,⁶ artifacts,^7,8 high cost, and inability to provide anatomical detail of difficult areas and very thin tissues.⁹

There is no doubt that interesting recent advancements have been made in TE. Nonetheless, further advancements are required to advance TE to its full potential via animal and clinical (human) trials. These animal and clinical trials, which may involve longitudinal (long-term) in vivo examinations of the same live in situ tissues, will make no tangible contribution to TE without an effective 3D biomedical imaging technique to monitor their success. An effective, noninvasive, 3D biomedical imaging technique to facilitate advancements should have the capacity to (1) track the process of cell growth into newly formed tissues (i.e., cell migration, cell-scaffold adhesion, and differentiation); (2) monitor scaffold biodegradation kinetics; and (3) examine the host-construct integration in a continued manner in the same animal or human without destroying or posing a risk to the animal or the tissues of interest.^10,11

Since the discovery of synchrotron light, different noninvasive 3D imaging techniques that are invaluable for biomedical applications and material sciences have emerged. The next section discusses the basic physics of major synchrotron-based biomedical imaging techniques with potential applications to bone and cartilage TE.

Synchrotron Imaging

Synchrotron light is an electromagnetic radiation produced when charged particles (i.e., electrons) are ejected from an electron gun by an electric field and then sped up in a linear accelerator. The particles are further accelerated to near the speed of light in a booster ring before being transferred to a storage ring. In the storage ring, bend magnets cause the electrons to change direction, and this results in a change in their velocity vector and, consequently, the radiation of synchrotron light. The properties of synchrotron light, to be discussed in detail next, significantly improve contrast sensitivity in X-ray imaging systems.¹² Due to the inherent advantages and potential of synchrotron-based light sources, several 3D biomedical imaging techniques have been developed.^13–18 The successful implementation of these methods in recent years demonstrates that synchrotron-based biomedical imaging techniques may be translated to TE applications. These methods will further advance TE by enabling noninvasive and longitudinal delineation of construct morphometry, neotissue growth, construct degradation kinetics, and host–implant interface dynamics.^8,10,11,17

Phase-contrast imaging

Phase-contrast imaging (PCI) uses variations in the phase shifts of emerging X-rays to characterize structural properties of samples with similar electron density, low atomic number materials, and soft tissues without the use of exogenous contrast agents.^19–21 At diagnostic X-ray energies, PCI relies on variations in the real part of the refractive index of tissues that are several orders of magnitude more important than the imaginary part used as the source of absorption contrast. In addition, unlike conventional diagnostic X-ray imaging that is dominated by incoherent Compton scattering, most implementations of PCI are based on coherent X-ray scattering for observation of refraction and interference.^19–22 Several PCI techniques have been developed based on obtainable information and differences in implementation. In the next section, we briefly discuss major PCI techniques that are currently or have the potential to be employed in TE.

Inline PCI

Inline PCI, also called propagation-based imaging or inline holography, had its major breakthrough in 1948. Inline PCI was the first phase-contrast technique pioneered; it requires no optical element (e.g., gratings or diffracting crystals) and has the simplest experimental setup.^23–26 It explores the phase shifts caused by variations in the refractive index and thickness of materials and captures these variations in measured intensity as edge enhancement between different regions.²⁵ This edge enhancement is a significant advantage of inline holography and results from the high lateral (spatial) coherence of inline PCI achievable using a small effective source or large object-to-detector distance.^23–26

An X-ray source such as a third-generation synchrotron X-ray source, a microfocus X-rays tube, or an ultrafast-laser-based plasma X-ray source is required to obtain high lateral coherence.^24,26,27 In addition, the object-detector distance is selected to fall in the Fresnel zone, a region between the absorption zone and the far field (Fraunhofer zone), as shown in Figure 1a.²⁸ The continuous phase variations result in enhanced and defined boundaries.^23–27 To optimize the sensitivity of this method, the object-detector distance should be limited to the Fresnel zone and completely avoid the Fraunhofer zone, and the spatial resolution of the detector, X-ray source size, and refractive properties of the samples should be considered.^23–27

FIG. 1.

Schematic diagrams illustrating the image acquisition setup of different phase-contrast X-ray imaging techniques: (a) inline PCI²⁸; (b) X-ray interferometry; (c) DEI²⁸; (d) DFI⁴¹; (e) differential PCI⁴⁴; and (f) coded-aperture phase contrast imaging.⁴⁹ DEI, diffraction-enhanced imaging; PCI, phase-contrast imaging; DFI, dark-field imaging. Color images available online at www.liebertpub.com/teb

X-ray interferometry

In 1966, Bonse and Hart pioneered the use of three Laue-case (LLL) analyzers in a method referred to as X-ray interferomerty.^13,29 X-ray interferometry is regarded as the most sensitive technique to measure phase shifts. It requires the high spatiotemporal coherence that is achievable by a high brilliance source such as a synchrotron or free-electron laser.^29–31 The set-up of the interferometer (Fig. 1b) employs three perfect diffracting crystals.³² The splitter crystal divides the incident X-ray beam into two coherent and spatially separated beams. One of the two beams passes through the sample, positioned between the mirror and analyzer, while the second beam acts as an unperturbed reference. The beams are then reflected via Laue-case diffraction at the mirror crystal. The two beams interfere at the entrance of the analyzer crystal, traverse the analyzer crystal, and produce intensity distributions detected by comparing the two beams. These intensity distributions are measured as phase shifts caused by the inhomogeneity in samples.^13,29–32

Diffraction-enhanced imaging

Diffraction-enhanced imaging (DEI), also known as analyzer-based imaging, was studied in the mid-1990s by Davis et al., Ingal and Beliaevskaya, and Chapman et al.^14,33,34 This method utilizes a three crystal setup that consists of a double crystal monochromator with asymmetric reflecting planes and an analyzer crystal (Fig. 1c).²⁸ The double crystal monochromator uses Bragg geometry to select a small energy band monochromatic beam from the incident polychromatic spectrum to traverse the object.^{12,14,33–35} The beam exiting the object then hits the analyzer crystal, which diffracts the X-rays that align only within its angular acceptance to the detector with a modulation given by the rocking curve (RC). All other photons that fall out of the angular acceptance window of the analyzer are of scatter origin, and their removal causes enhancement of the image contrast.^14,33,34 To obtain images, the analyzer is tuned at the half-maximum reflectivity of the RC and images are taken on both the high angle (θ_H) and low angle (θ_L) sides.^38–40 The intensities measured on the high angle (I_H) and low angle (I_L) sides are employed to calculate the refraction angle image (Δθ_Z; Eq. 1) and apparent absorption (I_R; Eq. 2) image obtained in the sensitivity direction Z, as follows: \documentclass{aastex}\usepackage{amsbsy}\usepackage{amsfonts}\usepackage{amssymb}\usepackage{bm}\usepackage{mathrsfs}\usepackage{pifont}\usepackage{stmaryrd}\usepackage{textcomp}\usepackage{portland, xspace}\usepackage{amsmath, amsxtra}\pagestyle{empty}\DeclareMathSizes{10}{9}{7}{6}\begin{document} \begin{align*}\Delta \theta_ {\rm Z} = \frac {{\rm I} _ {\rm H} {\rm R} (\theta_ {\rm L}) - {\rm I} _ {\rm L} {\rm R} (\theta_ {\rm H})} {{\rm I} _ {\rm L} \frac {{\rm dR}} {{\rm d} \theta} (\theta_ {\rm H}) - {\rm I} _ {\rm H} \frac {{\rm dR}} {{\rm d} \theta} (\theta_ {\rm L})} \tag {1} \end{align*} \end{document} \documentclass{aastex}\usepackage{amsbsy}\usepackage{amsfonts}\usepackage{amssymb}\usepackage{bm}\usepackage{mathrsfs}\usepackage{pifont}\usepackage{stmaryrd}\usepackage{textcomp}\usepackage{portland, xspace}\usepackage{amsmath, amsxtra}\pagestyle{empty}\DeclareMathSizes{10}{9}{7}{6}\begin{document} \begin{align*} {\rm I} _ {\rm R} = \frac {{\rm I} _ {\rm L} \frac {{\rm dR}} {{\rm d} \theta} (\theta_ {\rm H}) - {\rm I} _ {\rm H} \frac {{\rm dR}} {{\rm d} \theta} (\theta_ {\rm L})} {{\rm R} (\theta_ {\rm L}) \frac {{\rm dR}} {{\rm d} \theta} (\theta_ {\rm H}) - {\rm R} _ {\rm H} \frac {{\rm dR}} {{\rm d} \theta} (\theta_L)} , \tag {2} \end{align*} \end{document}

where R is the analyzer reflectivity as a function of analyzer rocking angle (θ), Δθ_Z is the refracted X-rays or the refraction angle intensity, and I_R is the apparent absorption intensity.^12,14

The apparent absorption image is then separated into absorption contrast and extinction contrast. The ability of DEI to separate these two contrasts is an advantage over conventional X-ray imaging methods.^12,14 The refraction effects of the X-rays and the thickness gradients of the sample greatly affect the refraction angle images. This results in intrinsic edge enhancement at the boundaries within tissues and enables DEI to provide clear and crisp refraction angle images. The absorption image relies on the linear integration of tissue absorption that, in turn, depends on the thickness of the sample.³⁶ Although DEI exhibits excellent scatter rejection, it is limited by its sensitivity to only the vertical components of the refracted beam; further research is required to extend its sensitivity.¹² DEI also experiences difficulty in characterizing homogeneous fine structures, and the images contain ultra-small-angle X-ray scattering (USAXS) properties that can degrade image contrast.^37–39 Other DEI-based methods are gradually evolving to address these challenges.

Rigon et al.³⁷ proposed an extension of DEI in a technique that employs the intensity measurements obtained from the peak and the two half -slopes of the RC. Unlike classical DEI, this technique also uses the second rather than the first derivative of the RC. This method generates apparent absorption, refraction, and USAXS images.³⁷ Multiple image radiography (MIR) is an optimized version of extended DEI that collects intensity measurements from five or more points on the RC. This gives MIR the capacity to overcome limitations associated with cross-talk. MIR uses the intensity measurements collected to also generate absorption, refraction, and USAXS images.³⁹

Analyzer-based phase contrast imaging (AB-PCI) is another DEI technique⁴⁰ in which the refraction-based images are taken only from one side of the analyzer RC, specifically at half maximum of the RC on either the high or low angle side. The slightly refracted X-ray beam exiting the sample is amplified by the perfect analyzer crystal placed between the object and the detector, thereby generating intensity variations that enhance the image contrast. This imaging setup does not give separate refraction images as in DEI; however, the technique has potential applications to live animal imaging due to its easy implementation, short imaging time, amplified contrast signal, and scatter rejection properties. This method has been reported to enable simultaneous visualization of both soft (cartilage and ligament) and hard (bone) tissues in situ with considerable structural detail.⁴⁰ This method may encounter limitations in sensitivity with regard to the wavefield in one direction. This limitation, which is characteristic of all analyzer-based phase-sensitive imaging techniques, may prevent the full delineation of samples with fibres oriented in many orientations, for example, cartilage. However, Li et al. suggested solutions that may be applied to this problem.⁴⁰

Dark-field imaging

Dark-field imaging (DFI) relies on monochromaticity and small angular divergence associated with its crystal optics. The Bragg-Bragg-Laue geometry of DFI simultaneously produces dark-field and bright-field images. The Bragg–Bragg geometry of the double crystal monochromator is used to generate a monochromatic beam from the incident beam, asymmetric diffraction of the beam, and reduction of the beam angular divergence.^15,41–43 The generated beam then hits and traverses the sample to be incident on the Laue-geometry analyzer, which then, with small-angle scattering, diffracts and splits the modulated beam as shown in Figure 1d.⁴¹ One of the two refracted beams produces intensities recorded by the area detectors as dark-field images, while the other beam generates bright-field images. This method requires no analyzer tuning and suppresses background illumination. Hence, DFI can be used to obtain high spatial resolution images of weak refraction-based signals exiting soft tissues.^15,41,42 DFI is also achievable using grating interferometry.⁴³

Differential PCI using X-ray gratings

While most PCI techniques require highly brilliant monochromatic sources, such as synchrotron sources, differential PCI (DPCI) works with low brilliance polychromatic X-ray sources.¹⁶ DPCI is also called Talbot-Lau interferometry. DPCI relies on angular filtering of the transmitted X-ray beam by microperiodic gratings followed by conversion of the angular beam deviations from refraction into intensity patterns recorded by the detector.^44,45 Depending on the brilliance of the X-ray source, DPCI requires two or three transmission gratings. Its intensity is separated into absorption, refraction, USAXS, and sometimes dark-field components. The setup (Fig. 1e) consists of the source grating or beam splitter (G₀), phase grating (G₁), and analyzer absorption (G₂) grating, with their respective periods (p) and distances from the detector.⁴⁴ G₀ consists of grids that divide the incident beam into an array of line sub-sources which are individually spatially coherent but mutually incoherent. The X-rays exiting each aperture of G₀ irradiate the sample and experience slight refraction. After the beam transmits the object, G₁ causes intensity modulation and splits the beam into two out-of-phase beams. One of these beams traverses the object and interferes with the other beam that exited G₀, in the region behind G₁ and in the plane of G₂, to produce periodic interference fringes. The linear periodic interference fringe patterns are transformed into intensity modulations that are recorded by the detector.^12,44,45 Unlike most PCI methods, DPCI enables a large field of view using large pixel-size detectors and requires less mechanical stability, which enables large image sampling.^16,45 In addition, DPCI is attracting clinical attenti/on because of its capacity to use low brilliance polychromatic X-ray sources to scan soft tissues at a high spatial resolution.^16,47

Coded-aperture phase-contrast imaging

Coded-aperture PCI (CAXPCI) is similar to DPCI; however, this method uses single or double noninterferometric gratings to generate a phase-contrast effect.^48–50 The major incident X-ray beam passes through “coded apertures” of the grating. The coded apertures that are placed just before the sample (the presample masks) diffract the incident X-rays and generate discrete, shaped, and smaller X-ray beams. These act as smaller X-ray sources and prevent inefficient radiation from traversing the sample (Fig. 1f).⁴⁹ X-rays from the presample coded apertures traverse the sample to the edge of the detector pixels, as defined by the postsample coded apertures or detector masks.^48–50 The detector and the detector mask are placed close to each other and aligned such that X-rays transmitted from the center of the detector masks align with the center of each pixel of the detector.⁵⁰ The sensitivity of the generated image is enhanced by illuminating only the edges of the detector pixels, as only a fraction of the photons admitted by the presample coded apertures reach the detector. This phenomenon is called the pixel edge illumination effect. The illuminated fraction and system sensitivity to phase contrast may be controlled by varying the alignment of the two sets of coded apertures.⁴⁸

The major advantage of CAXPCI is that it uses divergent, polychromatic, commercially available X-ray sources and area detectors. These advantages enable easy implementation in laboratory or clinical settings.^48–51 In addition, the method uses variable coded apertures to cause different edge illumination fractions and enables removal of unnecessary radiation dose via the presample coded apertures.⁴⁸ Despite the prospects of this method, it is practically difficult in a clinical diagnostic X-ray imaging system to increase the source-detector distance beyond 1 m, yet this is one of the major requirements for coherency.⁵¹ Furthermore, absorbed dose, exposure time, and image contrast trade-offs should be effectively monitored, as absorbed dose might increase in comparison with absorption-based methods.⁴⁸

K-edge subtraction

Contrast-enhanced K-edge subtraction (KES) imaging, also known as dichromography or dual-energy subtraction imaging, uses biocompatible and higher atomic number contrast agents to increase the sensitivity and contrast of low density tissues.⁵² It was pioneered by Jacobson in 1953 and uses either a synchrotron or compact X-ray tube source and a Bragg-case crystal monochromator system.^53–55 It relies on the large jump in attenuation coefficient (e.g., up to a factor of 6 for iodine) that occurs at the K-edge energy of the contrast agent, while the attenuation coefficient of the other components (e.g., surrounding matrix) of the object changes only minimally.^55–57 To implement KES imaging, two monoenergetic X-ray beams with defined energies of just below and above the K-edge of the contrast agent are selected from the incident polychromatic beam as shown in Fig. 2.¹² The sample is imaged with the two selected X-ray energies to simultaneously obtain two intensity images, one of which is subtracted from the other to give a difference image. The sensitivity of the difference image is directly proportional to the concentration of the contrast agent per image pixel.^12,52,54 The difference image (Eq. 3) is solved to generate the projected densities (images) of the matrix, ρ_Mt_M (Eq. 4), and the contrast material, ρ_ct_c(Eq. 5): \documentclass{aastex}\usepackage{amsbsy}\usepackage{amsfonts}\usepackage{amssymb}\usepackage{bm}\usepackage{mathrsfs}\usepackage{pifont}\usepackage{stmaryrd}\usepackage{textcomp}\usepackage{portland, xspace}\usepackage{amsmath, amsxtra}\pagestyle{empty}\DeclareMathSizes{10}{9}{7}{6}\begin{document} \begin{align*}\left[ \begin{matrix} {\rm r} _ {\rm L} \\ {\rm r} _ {\rm H} \end{matrix} \right] = \left[ \begin{matrix} - \ln \frac {{\rm I}_{\rm L}}{{\rm I} _ {\rm o}} \\ - \ln \frac{{\rm I}_{\rm H}}{{\rm I} _ {\rm o}} \end{matrix} \right] = \left[ \begin{matrix} \frac {\mu} {\rho_ {\rm M}} ({\rm E}_{\rm L}) \rho_{\rm M} {\rm t}_{\rm M} + \frac {\mu} {\rho_{\rm C}} ({\rm E} _{\rm L}) \rho_ {\rm C} {\rm t} _ {\rm C} \\ \frac {\mu} {\rho_ {\rm M}} ({\rm E} _ {\rm H}) \rho_ {\rm M} {\rm t} _ {\rm M} + \frac {\mu} {\rho_ {\rm C}} ({\rm E} _ {\rm H}) \rho_ {\rm C} {\rm t} _ {\rm C} \end{matrix} \right] \tag {3} \end{align*} \end{document} \documentclass{aastex}\usepackage{amsbsy}\usepackage{amsfonts}\usepackage{amssymb}\usepackage{bm}\usepackage{mathrsfs}\usepackage{pifont}\usepackage{stmaryrd}\usepackage{textcomp}\usepackage{portland, xspace}\usepackage{amsmath, amsxtra}\pagestyle{empty}\DeclareMathSizes{10}{9}{7}{6}\begin{document} \begin{align*}\rho_{\rm M}{\rm t}_{\rm M} = \frac{\frac{\mu} {\rho_{\rm C}} ({\rm E}_{\rm H}) {\rm r}_{\rm L} - \frac {\mu} {\rho_ {\rm C}} ({\rm E}_{\rm L}) {\rm r}_{\rm H}} {\frac {\mu} {\rho_ {\rm M}} ({\rm E}_{\rm L}) \frac {\mu} {\rho_ {\rm C}} ({\rm E}_{\rm H}) - \frac {\mu} {\rho_ {\rm C}} ({\rm E} _ {\rm L}) \frac {\mu}{\rho_ {\rm M}} ({\rm E} _ {\rm H})} \tag {4} \end{align*} \end{document} \documentclass{aastex}\usepackage{amsbsy}\usepackage{amsfonts}\usepackage{amssymb}\usepackage{bm}\usepackage{mathrsfs}\usepackage{pifont}\usepackage{stmaryrd}\usepackage{textcomp}\usepackage{portland, xspace}\usepackage{amsmath, amsxtra}\pagestyle{empty}\DeclareMathSizes{10}{9}{7}{6}\begin{document} \begin{align*}\rho_ {\rm C} {\rm t} _ {\rm C} = \frac{\frac{\mu} {\rho_ {\rm M}} ({\rm E} _ {\rm H}) {\rm r} _ {\rm L} - \frac {\mu}{\rho_{\rm M}} ({\rm E} _ {\rm L}) {\rm r}_{\rm H}} {\frac {\mu} {\rho _ {\rm M}} ({\rm E} _ {\rm L}) \frac {\mu} {\rho_ {\rm C}} ({\rm E} _ {\rm H}) - \frac {\mu} {\rho_ {\rm C}} ({\rm E} _ {\rm L}) \frac {\mu} {\rho_ {\rm M}} ({\rm E} _ {\rm H}) ,} \tag {5} \end{align*} \end{document}

FIG. 2.

Schematic diagram of KES imaging setup with a single bent Laue monochromator with splitter.¹² KES, K-edge subtraction. Color images available online at www.liebertpub.com/teb

where I is the photon intensity; subscripts M, C, L, and H indicate the matrix, the contrast material, below K-edge, and above K-edge, respectively; ρ is the density of the material; t is the thickness of the material; I₀ is the number of photons in the incident beam; \documentclass{aastex}\usepackage{amsbsy}\usepackage{amsfonts}\usepackage{amssymb}\usepackage{bm}\usepackage{mathrsfs}\usepackage{pifont}\usepackage{stmaryrd}\usepackage{textcomp}\usepackage{portland, xspace}\usepackage{amsmath, amsxtra}\pagestyle{empty}\DeclareMathSizes{10}{9}{7}{6}\begin{document} $$\frac {\mu} {\rho} ({\rm E})$$ \end{document} is the mass attenuation coefficient of the material; and r= \documentclass{aastex}\usepackage{amsbsy}\usepackage{amsfonts}\usepackage{amssymb}\usepackage{bm}\usepackage{mathrsfs}\usepackage{pifont}\usepackage{stmaryrd}\usepackage{textcomp}\usepackage{portland, xspace}\usepackage{amsmath, amsxtra}\pagestyle{empty}\DeclareMathSizes{10}{9}{7}{6}\begin{document} $$\ln {\frac {\rm I} {\rm I} _0}$$ \end{document} .^12,52,54

The presence of bones, motional image blurring, and cross-over artifacts caused by beam crossing in addition to the technical complexity and rarity of synchrotron sources constitute major obstacles to KES imaging.⁵⁶ Thus, new applications are being developed to optimize the method. KES has been combined with CT (KES-CT) for extensive quantification of the spatial distribution of contrast agent in samples.⁵⁷ In addition, Zhong et al. investigated a lab-based rotating anode X-ray source and single bent Laue monochromator for KES angiography to address the scarcity of synchrotron facilities.⁵⁵ The recent development of three-energy KES has also contributed to advancements in KES imaging. In three-energy KES, a third image is simultaneously taken using a harmonic energy far above the K-edge energy of the contrast agent. The third image resolves the bone component, which is an artifact source in the images acquired above and below the K-edge.^56,58 Similarly, near-edge spectral imaging (NESI), an extended version of KES imaging that enables scanning from an energy just below to an energy just above the K-edge of the contrast material, has been developed. NESI has numerous advantages over KES imaging, including increased photon flux (up to a factor of 5), elimination of motion artifacts, and easy characterization of additional components such as bone.⁵⁶

Furthermore, shaping the X-ray spectrum such that the two energies are very close to the K-edge of the contrast material has been reported to optimize the sensitivity of KES.⁵⁹ In addition, the use of a bent Laue monochromator has been reported to optimize photon flux, integrated reflectivity, and tolerance to diffraction due to a large solid diffraction angle.⁵⁶ Finally, the cytotoxicity of contrast agents and the selection of a contrast material with a suitable absorption energy to prevent photon starvation during biomedical imaging should also be taken into consideration.⁶⁰

Synchrotron radiation microcomputed tomography

Synchrotron radiation microcomputed tomography (SR-μCT) is the synchrotron-based version of the μCT systems introduced by Feldkamp et al.⁶¹ The high photon flux (typically about 10⁸ photons/s), high brilliance of the X-ray source, small angular beam divergence and consequent improved spatial resolution and higher signal-to-noise ratio, and faster acquisition time are benefits of SR-μCT over standard μCT systems. In addition, the ability to tune the synchrotron X-ray energy addresses issues with regard to beam hardening artifacts, reconstruction artifacts, and geometrical artifacts that occur in standard μCT imaging.^63–66 SR-μCT employs a collimated monochromatic beam generated from a double crystal monochromator for generation of absorption contrast. The monochromatic beam illuminates a sample mounted and fixed on a rotatable stage as illustrated in Fig. 3.⁶⁹ The intensity of the transmitted beam is recorded by an electro-optic detector positioned close to the sample to avoid phase contrast.^61–66

FIG. 3.

Schematic diagram of part of a synchrotron radiation μCT imaging setup. Inset images illustrate the effect of X-ray propagation distance on inline phase contrast images (transition of absorption-contrast image to phase-contrast image by sample-to-detector distance).⁶⁹ μCT, microcomputed tomography. Color images available online at www.liebertpub.com/teb

An extension of SR-μCT is the use of synchrotron-based phase contrast techniques, as discussed earlier, in what is known as phase-contrast microcomputed tomography (PC-μCT).^64–67 PC-μCT combines the properties of SR-μCT and phase-contrast techniques for simultaneous generation of phase contrast alongside the absorption contrast usually produced in SR-μCT. This technique is much more sensitive, especially to soft tissues, and can generate absorption images at greatly reduced absorbed doses; this method is more prevalent in bone and cartilage TE characterization.^64–67

Applications of Synchrotron Imaging to Bone and Cartilage TE

Potential applications to bone TE

Bone tissues have high absorption coefficients, therefore making their imaging easier than for soft tissues. Several conventional noninvasive imaging methods, such as μCT, ultrasound, and MRI, are commonly used for bone imaging and bone TE (BTE) applications.^1,3,4 BTE requires delineation of surrounding soft tissues, knowledge of the degradation kinetics of the implanted construct, and bone remodeling profiles alongside host bone. As a result, imaging techniques with the capacity to simultaneously visualize both hard and soft materials with high and low absorption coefficients are in high demand. Synchrotron-based imaging methods provide information and possess different advantages above and beyond conventional imaging techniques. A summary of the advantages and disadvantages of synchrotron-based imaging techniques for TE applications is provided in Table 1. This information includes architectural properties, degradation kinetics of constructs, regenerated tissue quality as characterized by vascularization, and integration of the constructs with the surrounding host tissue. In the next section, we discuss the current status and capabilities of synchrotron-based techniques used in BTE and conditions similar to those encountered in BTE.

Table 1.

Advantages and Disadvantages of Commonly Used Synchrotron X-ray Imaging Techniques

	Advantages	Disadvantages
SR-μCT	Simple experimental setup	Cannot specifically identify tissue types⁶⁶
	Image pixel size of approximately 0.28 μm⁶⁶	High absorbed radiation dose due to high-resolution imaging
	Predominantly used for ex vivo and in vitro imaging of hard tissues, e.g., bone imaging^65–76
Inline PCI	Simplest phase-sensitive imaging experimental setup	High energy, lower dose imaging; scatter production at high energy
	Edge enhancement for weakly absorbing materials, e.g., soft tissues^23–26	Practical limitations with extreme propagation distances; infrastructure limitation
Inline PC-μCT	Specific identification of tissue types⁶⁶	High absorbed radiation dose due to high resolution imaging; limitation for in vivo imaging
	Simple experimental setup⁷⁶
	Image pixel size of approximately 0.7 μm⁷³	Heavy image data processing; requires data-intensive computing systems
	Excellent for soft and hard tissue imaging ex vivo and in vitro^65,68,76
DEI	Acquisition of absorption- and refraction-based contrast sensitive to small density change^77–81	Does not compensate for USAXS; sensitive to only the vertical components of the refracted beam^37–39
	Scatter free; enables higher energy, lower dose imaging	High sensitivity of the setup to motion, e.g., artifact production by object movement or analyze sensitivity to background physical noise
	Suitable for soft and hard tissue imaging^77–81	Requires at least two sets of scans; longer imaging time, patient discomfort
	Suitable for thick object imaging in vivo, ex vivo, and in vitro^77–81
DPCI	Large field of view (large sample) imaging	Limited angular acceptance of the beam; grating apertures are longer in the X-ray propagation direction than in the vertical direction
	Acquisition of differential phase, USAXS, and absorption contrasts⁴⁷	Sensitive to vibration as aperture dimensions are very small
	Use of low-brilliance polychromatic X-ray sources^16,45	Field of view is limited by grating's manufacturing process⁴⁷
CAXPCI	Use divergent, polychromatic, and commercially available X-ray sources and area detectors; easier transfer to clinical application	Difficulties with increasing the source-to-detector distance beyond 1 m for clinical systems; challenges with providing the major beam requirement of coherency
	Unwanted beams are removed before the beam passes through the sample; lower radiation dose imaging^48–50
KES-CT	Quantitative and qualitative imaging of very low absorbing or refracting objects (i.e., cells)	Reliance on contrast materials; exogenous and risk of cytotoxicity
	Identification and quantitative separation of different components of the object, e.g., contrast material, bone, and soft tissues⁶⁰	Requires two sets of scans; longer imaging time, patient discomfort
DFI	No background illumination	High radiation dose than inline PCI and absorption imaging⁸²
	Detection of differences in microscopic density fluctuations in soft tissues
	Suitable for soft tissue imaging^15,82,83

CAXPCI, coded-aperture phase-contrast imaging; DEI, diffraction-enhanced imaging; DFI, dark field imaging; DPCI, differential phase-contrast imaging; KES-CT, K-edge subtraction computed tomography; PC-μCT, phase contrast microcomputed tomography; SR-μCT, synchrotron radiation microcomputed tomography; USAXS, ultra-small-angle X-ray scattering.

Synchrotron radiation microcomputed tomography

SR-μCT is an absorption-based imaging technique that can be used for 3D quantitative and qualitative characterization of tissue properties at a spatial resolution even smaller than 1 μm.⁶⁸ It exhibits faster scanning than laboratory μCT⁶⁵ and benefits from the advantages of synchrotron source X-rays.^66,67 SR-μCT has a demonstrated capacity to obtain 3D information about trabecular and cortical bone microarchitecture at a spatial resolution less than 10 μm.^63,65 SR-μCT has also been employed for 3D characterization of scaffold properties, such as pore connectivity, pore size, wall thickness, and anisotropy, as well as functional changes resulting from tissue regeneration.^2,69 In fact, SR-μCT enables characterization of scaffold micro- and macro-structures in a manner not possible with two-dimensional (2D) scanning electron microscope (SEM) analysis.⁶⁷ SR-μCT has been employed for bone imaging and scaffold characterization as well as in a few BTE studies. SR-μCT was used by Yue et al.⁶⁶ to identify scaffold morphology, mineral distribution within scaffold pores, and tissue ingrowth in 4-week-old explants of a bioactive glass foam scaffold implanted between the muscle and tibia of a mouse (Fig. 4).⁶⁶ In a similar study,⁷⁰ SR-μCT was used to successfully identify scaffold architecture and bone ingrowth into cell-loaded hydroxyapatite scaffolds implanted in immunodeficient mice for 8 weeks. The bone ingrowth was estimated in terms of total volume fraction, distribution, and thickness of the newly formed bone tissue in the pores of the implant and the scaffold architecture in terms of porosity and spatial distribution of walls.⁷⁰ The same group explored the ability of SR-μCT to examine the progressive resorption of and bone ingrowth into scaffolds implanted in immunodeficient mice for longer repair times of 8, 16, or 24 weeks.^71,72

FIG. 4.

SR-μCT images of a bioactive scaffold: (a) scaffold morphology and pores before implantation; (b) 4 weeks after implantation (ex vivo) between the muscle and tibia of a mouse showing changes in the scaffold and some tissue growing into the pores of the scaffold; and (c) calcium distribution within the scaffold.⁶⁶ SR-μCT, synchrotron radiation microcomputed tomography. Color images available online at www.liebertpub.com/teb

The use of gold nanoparticles (GNPs) as contrast agents for enhancing X-ray attenuation in the region of interest using SR-μCT has also been investigated. In two recent studies, Astolfo et al. used SR-μCT to localize and track GNP-labeled cells.^73,74 In one of these studies,⁷³ they investigated the increased sensitivity obtained from inclusion of GNPs ex vivo, by injecting GNP-labeled gliomal cells into the brain of adult male Wistar rat for 16 days, and in vitro, by suspending the labeled cells in 5% agarose gel. The in vitro experiment demonstrated that the enhanced contrast enabled assessment of individual GNP-loaded cells along with the surrounding lacunae. Although the in vivo imaging resolution was four times lower and the number of projections was reduced compared with similar ex vivo imaging (to address the dose issue), it was still possible to localize the clusters of cells (Fig. 5). The latter study⁷⁴ showed a reasonable compromise between dose and image resolution, which is the case in most longitudinal in vivo studies. Although the application of SR-μCT to track the performance of labeled cells in TE constructs is yet to be investigated, successful results^73,74 may be extrapolated to BTE for tracking tissue ingrowth into TE constructs.

FIG. 5.

Gold nanoparticle-enhanced 3D rendered SR-μCT images showing the effect of varying spatial resolution: (A) low dose in vivo image and (B) high dose ex vivo image.⁷⁴ 3D, three-dimensional. Color images available online at www.liebertpub.com/teb

Several studies have applied SR-μCT with great success to bone tissue imaging and the characterization of the functionality of scaffolds used in BTE.^65–75 However, the majority of these examinations focused on tissue samples or in vitro studies. This is mainly due to the high radiation dose coupled large exposure times to achieve very high resolution. This trade-off should be satisfied when high quality and quantitative imaging is desired. The inclusion of nanoparticles as contrast agents may enhance tissue sensitivity even when spatial resolution is reduced.^73,74 Overall, SR-μCT is a suitable imaging technique for highly detailed (e.g., cellular level) qualitative and quantitative nondestructive characterization of constructs and tissue regrowth. However, due to the limitations associated with this technique, such as small sample size (less than 1 mm³), high radiation dose, and scanning time, candidate samples may be limited to excised specimens and small animals (noninvasive in vivo imaging).⁷⁵

Inline PC-CT (and PC-μCT)

Even though applications of SR-μCT have been very successful with regard to the characterization of scaffolds, bone, and bone growth kinetics,^65–75 this technique is unable to clearly identify differences between several low density tissues.⁶⁶ PC-μCT can be used to obtain phase-contrast images alongside absorption contrast images of tissues. On one hand, its phase contrast provides sensitivity to poorly absorbing biological samples with low absorption contrast, such as soft tissues, such that it can differentiate various tissue types. On the other hand, its absorption contrast provides greater detail for tissues with high absorption contrast.^8,65 Several groups have used PC-μCT or inline PC-CT to characterize bone tissues and the functionality of biomaterial scaffolds or constructs used in BTE. In one such study, inline PC-μCT was used to identify different phases resulting from the organization of the extracellular matrix, from fibrils into networks, formed within scaffolds seeded with stem cells after 15 days in culture.⁶⁸ Weiss et al.⁷⁵ used PC-μCT to examine the microstructure, bone ingrowth, mineralized and unmineralized bone tissue network, and tissue–implant interface of an injectable bone substitute and macroporous blocks implanted in a rabbit model and reported results consistent with conventional 2D histomorphometric analysis.⁷⁵ Similarly, Sun et al.⁸⁴ successfully applied inline PC-CT and DEI to characterize the repair of osteonecrosis in rabbits using nano-hydroxyapatite/collagen and autologous mesenchymal stem cells. Both techniques demonstrated identification of the biomaterial–host interface, bone tissue formation, and substitution of biomaterials with newly grown tissues over a period of 12 weeks.⁸⁴ In another study,⁷⁶ SR-μCT and PC-μCT were investigated and compared with regard to imaging bone ingrowth coupled with angio- and microvasculogenesis. Although scaffold materials and mineralized bone were visible using both imaging techniques, PC-μCT provided information on the scaffold material, bone ingrowth, and vessel network resulting from microvasculogenesis (Fig. 6).⁷⁶

FIG. 6.

(A) Three-dimensional SR-μCT image showing scaffold material as white and bone ingrowth as brown 24 weeks postimplantation in immunodeficient mouse; (B) 3D PC-μCT image showing scaffold material in white, bone ingrowth in brown, and vessel ingrowth in green at 24 weeks postimplantation in immunodeficient mice; and (C) magnified image of identified region in (B).⁷⁹ PC-μCT, phase contrast microcomputed tomography. Color images available online at www.liebertpub.com/teb

During the bone remodeling stage of a tissue-engineered joint, certain details with low or similar densities are challenging to decipher using absorption-based imaging techniques, even at an optimum photon energy.⁶⁶ Phase contrast-based imaging is a potential technique for deciphering structural details and microvasculogenesis during tissue regeneration and remodeling. In addition to regeneration progress, inline PC-μCT can decipher scaffold biodegradation kinetics and other nonbone surrounding tissues.^68,75,76 Although SR-μCT is currently the most prevalent technique used for bone imaging and BTE, inline PC-μCT's experimental simplicity, contrast resolution that is many orders of magnitude higher than the absorption contrast, and ability to decipher the growth of vessel networks resulting from angio- and microvasculogenesis, changes in scaffold material, and bone ingrowth may be attractive for BTE applications.^65,68,75,76

PC-μCT may be an ideal technique for highly detailed (e.g., cellular level) qualitative and quantitative characterization of engineered constructs, hard and soft tissues regrowth in vitro, and tissue samples obtained from small animals. However, extrapolating this method for use in in vivo studies requires that certain trade-offs be made regarding image quality and absorbed radiation dose. With appropriate manipulations, inline PC-CT may be a better candidate than SR-μCT for the characterization of tissue-engineered repair in vivo in large animals.⁸⁴

Diffraction-enhanced imaging

Excellent scatter rejection and the ability to separate apparent absorption contrast (vs. extinction contrast) and refraction contrast are advantages of DEI over conventional X-ray imaging methods. As such, researchers have used this method for imaging bone,⁹³ the bone–cartilage interface,⁹⁴ and BTE applications.⁷⁰ Cooper et al.⁹³ successfully used DEI to characterize trabecular bone architecture in human cadaveric radii, forearm, and hand within intact tissues. DEI provided information about boundaries between bone tissues with differing refractive indices: inter-individual differences in trabecular texture, cortical pores, and resorption spaces associated with newly forming osteons. They concluded that DEI may be suitable for the detection of early changes associated with bone loss.⁹³ Connor et al.^95,96 report increased contrast for imaging interfaces and contrast-to-noise-ratio gain using DEI compared with SR radiography in their study of interface gaps when titanium pins were implanted into bone defects. Similarly, DEI and inline PC-CT used as complementary techniques enabled the characterization of osteonecrosis repair.⁸⁴ In fact, the image details showed the boundary of the bone graft and the trabecular network (Fig. 7).⁸⁴ Although DEI is not a currently prominent characterization technique in BTE, refraction images obtained alongside the apparent absorption images from DEI can provide additional information about the functionality of the constructs and the quality of their integration with host tissues.²²

FIG. 7.

Visualization of tissue engineered osteonecrosis repair using inline PC (A, B) and DEI (C, D) as complementary imaging techniques; bone defect repair at 8 weeks (A, C) and 12 weeks (B, D) postoperation.⁸⁴ Black arrows show the boundary of the bone graft area with surrounding tissue. Color images available online at www.liebertpub.com/teb

K-edge subtraction/K-edge subtraction computed tomography

Researchers have evaluated KES for applications to mammography, coronary angiography, and cancer cell imaging.^{53,54,58–60} However, the potential of KES (or KES-CT) for TE applications has not been considered. The ability of KES to differentiate between materials with varying attenuation coefficients and isolate these different attenuations makes it attractive for quantitative imaging of bone tissue-engineered scaffolds, possibly new tissue ingrowth, and surrounding bone tissue. Though the method is yet to be investigated for any TE application, including bone, Cooper et al. recently investigated the principle in a rat specimen treated with strontium (Sr) ranelate.⁶⁰ They reported that they were able to trace Sr uptake within the microarchitectural features of vertebrae (Fig. 8). Comparison of the Sr map obtained by KES-CT and electron probe microanalysis showed a slightly lower concentration of Sr in the KES-CT (1.36% to ∼2% by mass) data (Fig. 8). However, KES-CT is nondestructive and may be preferable with regard to preserving the sample and enabling mapping of the bone formed over the dosing period.⁶⁰ Although this study was not directed toward BTE, it demonstrates the potential of this method for tracking and enhancing bone growth in BTE through the use of Sr as a contrast agent and North American nutraceutical Sr supplements for osteoporosis.⁶⁰

FIG. 8.

(A) Volumetric 3D KES images showing the uptake of Sr and the trabecular, cortical, and periosteal bone remodeling/growth; (B) magnified image of identified region used for comparison of the distribution of Sr by KES and electron probe microanalysis (EPMA).⁶⁰ Sr, strontium. Color images available online at www.liebertpub.com/teb

In summary, studies to date show that different SR imaging techniques offer great potential for BTE applications. They can aid in making advances in current TE studies as well as future extensions of TE studies to longitudinal in vivo animal and clinical studies. For live animal imaging, currently available methods should be improved to reveal detailed information about bone ingrowth, scaffold degradation kinetics, and interaction with surrounding tissues at a lower radiation dose. Ex vivo BTE imaging would be highly valuable for optimization of scaffold properties and design to better repair bone damage. However, determining the appropriate trade-offs to positively contribute to the advancement of these techniques toward safe and long-term characterization in BTE for in vivo studies is crucial.

Potential applications to cartilage TE

The high water content and low density of cartilage tissue produces negligible X-ray attenuations that limit its visualization using conventional absorption-based X-ray imaging techniques.^6,68 MRI is commonly used to examine cartilage and to identify damage and loss of tissue without contrast media.⁶ Nevertheless, the poor spatial resolution of MRI and inability to resolve specific tissue types when compared with synchrotron-based techniques is the bottleneck of this modality.^6,34 Synchrotron imaging techniques were initially investigated due to the demand for superior imaging techniques for early and accurate diagnosis of cartilage damage and disease, such as the detection of osteoarthritis (OA) in its early stages. As such, most studies in the literature have concentrated on synchrotron imaging techniques for the visualization and characterization of healthy and damaged/diseased cartilage tissue as well as the cartilage–bone interface. A review of these studies, in addition to the few studies that actually investigated techniques for cartilage TE, will inform and provide a better understanding with regard to the potential and limitations of these techniques for cartilage TE applications.

Inline PCI (and PC-μCT)

The simplicity of implementation, provision of outstanding contrast of less dense samples without the use of contrast agents, and edge enhancement are notable advantages of inline PCI. Inline PCI is a powerful tool for characterization of the cell-scaffold matrix, new tissue ingrowth, and bone–cartilage interface where the refractive indices vary greatly.^6,25,28 The combination of inline PCI with either CT or μCT has been studied for 3D visualization of articular cartilage.^6,17 One such study, conducted by Ismail and colleagues, used a bench-top microfocus X-ray source with polychromatic, incoherent X-rays as well as synchrotron coherent X-rays to examine the cartilage–bone interface and different zones of cartilage.⁶ They recorded edge enhancement, especially using synchrotron X-rays, that was sufficient to visualize low density cartilage and the cartilage–bone interface with transitional zones from the articulating surface down to subchondral bone.⁶ In a similar study, Zehbe et al.⁸⁵ used PC-μCT for 3D qualitative and quantitative characterization of articular cartilage; the 3D rendered images revealed information which was superior to that obtained from conventional serial histology (Fig. 9). Using the 3D images, various tissues under physiological and pathological conditions were differentiated. In addition, tissue structure and cellular level changes, such as spatial cell density and the shape and orientation of the lacunae inside the soft tissues, were quantified without destruction of the tissues.⁸⁵ Using high-resolution PCI, Choi et al.⁸⁹ characterized the microstructural features of healthy cartilage as well as inflammatory and pathological changes in the arthritic joint of mice in vivo, also showing results comparable to those of microCT or histological analysis.

FIG. 9.

Images showing cellular information from the frontal lacuna in the soft tissue of an articular cartilage: (a) original scanning electron microscope image from histological tissue; (b) region of interest containing lacuna doublet with chondrocytes in the centres from histological tissue; (c) similar sliced image data from inline PC-μCT; (d) 3D rendering of (c); and (e) quantitative analysis of cell density from inline PC-μCT (adapted from Zehbe et al.⁸⁵). Color images available online at www.liebertpub.com/teb

Synchrotron radiation microcomputed tomography

Tissue-engineered scaffolds have also been characterized using SR X-ray imaging techniques.^{85–87,97,98} In studies using SR-μCT, a quantitative description and information with regard to the scaffold microstructure and/or cell clusters in culture or in 3D scaffolds were obtained. Due to the light element constituents of cells (water), contrast agent staining, that is, Au-lysine and silver enhancer, were used to increase the contrast between cells and the surrounding medium in absorption-based imaging.⁹⁸ In a similar manner, using SR-μCT yielded good results with regard to the visualization of chondrocytes embedded in a porous gelatin scaffold and physical properties of the scaffold including porosity, surface area, circulation, and pore directionality (width and height of pores).⁸⁸ Although most reported studies concentrate on in vitro examinations, the techniques and ideas used may be extrapolated to the visualization of tissue-engineered repair in exercised tissue samples and future in vivo applications. Similar to BTE, SR-μCT is more applicable to high resolution, detailed ex vivo investigations of tissue-engineered scaffolds and constructs, which are used to evaluate excised scaffolds or tissue constructs for improving TE strategies.

X-ray interferometry

X-ray interferometry is commonly used for characterizing variations in microcalcifications, fat, blood vessels, and soft tissues. It has been reported to be beneficial for the early detection of breast cancer, colon cancer, necrosis, and kidney disease.^13,29 X-ray interferometry is suitable for deciphering minute variations in densities within less dense soft tissues such as cartilage.^77,108 While little is known about the application of X-ray interferometry to cartilage TE, this method has great potential for TE applications, especially for the delineation of variations in soft tissues resulting from new tissue growth. However, the heat radiated from the body can deform the interferometer, as samples are typically placed close to the crystal lamellae.²⁸ In addition, this method may be unsuitable for resolving objects with sharp soft-hard tissue structural boundaries with a large refractive index difference, such as the bone–cartilage interface or implant–host interface.²⁸ The good news is that researchers have continued to explore the method and have recently reported some progress.¹⁰⁸

Diffraction-enhanced imaging

DEI is capable of scatter rejection, is sensitive to density differences in tissues, can be optimized through varying the analyzer angular setting without increasing the radiation dose, and provides multiple types of information such as absorption, refraction, and extinction images.⁹⁹ Hence, this method is promising for the 3D characterization of microstructural properties of regions with varying densities, especially in low X-ray absorbing materials such as native and tissue-engineered cartilage. DEI has been used to visualize cartilage tissue with appreciable structural detail in both ex vivo^77–79 and in situ samples.^17,80,81 High levels of detail in the cartilage structure, such as the structural organization of collagen fiber bundles within the articular cartilage, have been revealed using a DEI system.⁹⁸ Issever et al. reported the visualization of cartilage matrix, including changes in hypodensities that were strongly hypothesized to be chondrocyte lacunes.⁷⁷ Similarly, Wagner et al. used color-coding DEI to explore the internal structure of healthy and pathological joints. In their study, a comparison of DEI with MRI showed that the same level of structural detail may be revealed by MRI but only after a much prolonged exposure which might not be practical with a patient.⁷⁹ Muehleman et al.⁸⁰ differentiated stages of a cartilage lesion from a normal state to a down-to-bone erosion state in intact canine joints in situ using a DEI system. The high spatial resolution combined with the refraction-based mechanism of DEI enabled visualization of early-stage cartilage degeneration and defects in intact human knee and ankle joints in situ.⁸⁰ These results were further confirmed by histological and gross analysis.⁸¹ AB-PCI was also used for the characterization of osteoarthritic and normal cartilage matrices both ex vivo and in vivo.^40,91,92 In fact, the investigated technique enabled quantitative and qualitative characterization of the zonal pattern in the cartilage matrix, zonal thicknesses, chondrocyte homogeneity and alignment, and matrix fibrillation (Fig. 10B, C).⁹¹ Coan et al.⁹² tested this imaging modality on in vivo samples and observed a high level of contrast for depicting anatomic structural details and pathological features of an osteoarthritic articular joint.⁹² Similarly, Li et al. imaged intact human knee and obtained structural details such as cartilage tissue, cruciate ligaments, loose connective tissue, menisci, and chondrocalcinosis (Fig. 10A).⁴⁰

FIG. 10.

(A) AB imaging of cadaveric human knee joint showing enhanced contrast imaging of soft tissues and bone with structural details in situ⁴⁰ and (B) AB-CT imaging of cartilage specimen ex vivo, with zonal structural detail comparable to (C) histological analysis.⁹¹ Arrows in B, C are pointing at the same locations in both images to compare visibility of minute details. AB-CT, analyzer-based computed tomography. Color images available online at www.liebertpub.com/teb

Researchers have conducted extensive studies in cartilage imaging and diagnosis of cartilage disease using DEI and ABI. However, very few studies have explored the potential of DEI and ABI for characterization in cartilage TE. DEI-CT has proven to be ideal for visualizing thick TE sections and samples with varying densities and tissue types, such as scaffolds, OA affected cartilage, and newly regenerated tissue. In a recent study, Izadifar et al. characterized TE scaffolds implanted in the knee cartilage of a piglet joint using DEI-CT, inline PC-CT, and MRI.¹⁷ The results demonstrate the superiority of DEI-CT over inline PC-CT and MRI (Fig. 11A–C) for imaging TE scaffolds in situ. The results also show that DEI-CT could effectively delineate the cartilage microstructures and track the scaffolds and different soft tissues surrounding the joint better than the other two methods (Fig. 11D).¹⁷ In a similar study by Zhu et al., DEI was compared with laboratory-based radiography, SR-radiography, and inline PCI at the same energy.¹⁸ Their results show that DEI offered better structural and microstructural quantification of soft tissues over the other three methods because of its ability to reject X-ray scatter.¹⁸ DEI and ABI are not only capable of providing information about cartilage, bone, and scaffolds present in the joints, but may also offer supportive information about surrounding tissues such as tendons, ligaments, adipose pads, and skin.^17,40

FIG. 11.

Images of scaffolds implanted in the lateral femoral cartilage of a piglet stifle joint, with arrows showing the position of the scaffolds: (A) inline phase-contrast CT acquired at 37 μm; (B) DEI-CT acquired at 37 μm; and (C) magnetic resonance imaging acquired at 310 μm. (D) Tissues visible in the DEI images.¹⁷

DEI image quality may be enhanced without changing the resolution but rather by changing positions on the RC. In addition, DEI produces absorption, extinction, and refraction contrasts that translate into more information than what is obtainable from normal absorption-based imaging. Thus, DEI can facilitate noninvasive in vivo studies of 3D tissue-engineered constructs, cartilage ingrowth into the constructs, and surrounding hard and soft tissues in living animal models, and thus demonstrates potential for future clinical studies.

Dark-field imaging

DFI with an asymmetric-cut monochromator with Bragg geometry and a Laue-case analyzer crystal has also been investigated for cartilage visualization. The absence of background illumination caused by nonrefracted X-rays, its simplicity when compared with DEI, and its single exposure imaging capability enables DFI to be used to obtain higher refraction-based image contrast and minute details of soft tissues such as cartilages.^{15,41–43,82,83} Ando et al.⁸² examined the morphology of articular cartilage at a femoral head and a shoulder in a human cadaver under simulated clinical imaging conditions using DFI and ordinary X-ray absorption imaging. Their results (Fig. 12a, b) obtained using DFI clearly depict the cartilage region better than an ordinary absorption X-ray image. Interestingly, they concluded that DFI possesses clinical potential for accurate assessment of articular cartilage and associated disorders.⁸² Despite little being known with regard to the use of DFI for TE applications, its single exposure imaging principle, clinical potential, and ability to simultaneously image both cartilage and subchondral bone demonstrate promise for TE applications involving many tissue types.

FIG. 12.

(a) DFI image clearly showing the articular cartilage of a human femur joint immersed in water to mimic clinical conditions; (b) ordinary X-ray absorption image of the same femur joint; and (c) image of the distal end of a human femur immersed in formalin clearly showing the articular cartilage and femoral condyle.⁸²

Useful imaging specifications and parameters employed in the earlier studies for imaging cartilage and bone tissues and/or engineered constructs are summarized in Table 2 for interested readers.

Table 2.

Summary of Imaging Parameters Used for Synchrotron Radiation-Based Imaging of Engineered Constructs or Bone and Cartilage Tissues

Technique	X-ray source- imaging energy	Sample condition	Image resolution (μm)	Field of view	Exposure time	Dose rate	Reference	Segmentation technique/3D rendering
SR-μCT	Synchrotron; 9–27 keV	Excised—fresh/freeze dried	1.4–5	2–10 mm	0.7–3 s		66,67,70,74,85 –88	Histogram-, Gaussian-, intensity-, threshold-based segmentation
								3D rendering: OsiriX v. 3.0, VGStudio MAX v. 1.2.1, AMIRA 5.2, ImageJ
Inline PCI	Benchtop (X-ray tube); 70 kVp	Excised, thin sections (∼500 μm)	13	—	1 min	—	6
Inline PCI	Synchrotron; 7–40 keV	Excised—fresh/in vivo	1.48–37	1.18–4×0.89–74 mm²	25 ms–2.5 s	1.6–3 Gy/s	6,17,89	Histogram-based segmentation
						0.8–3 Gy (total)		3D rendering: Avizo
DEI-CT	Synchrotron; 17–50 keV	Excised—fresh/fixed	3.6–50	4 mm (vertical)×74 mm	3.5–7 s	0.01–0.34 mGy/s	17,77 –79,81,90	Histogram-based segmentation
						20 mGy–0.2 Gy (total)		3D rendering: Avizo
AB-PCI	Synchrotron; 26–52 keV	Excised—fixed/in vivo	8–47	9×9 cm²	0.5 s-planar imaging	0.3 mGy/s	91,92
					67 ms/projection
Inline PC-μCT	Synchrotron; 5–34 keV	Fixed	0.7–10		1–10 s	0.2 Gy (total)	25,68,75,76,85	Thresholding-based segmentation
								3D rendering: VGStudio MAX 1.1

AB-PCI, analyzer-based phase-contrast imaging; 3D, three-dimensional.

Research Issues and Future Directions

The inherent advantages of a synchrotron light source help overcome numerous intrinsic limitations of conventional imaging systems, such as poor spatial resolution, soft tissue contrast deficiency, X-ray filtration, high absorbed dose, several modality-specific artifacts, and so on.^7,64 Enhancing the capability of SR-based imaging techniques to effectively visualize scaffolds designed to aid tissue regeneration ex vivo and to enable progressive monitoring of the associated growth and construct degradation kinetics in living animals is a necessity. In the next section, we discuss issues and areas that require further development to make SR-based techniques effective for TE applications.

Contrast media as sensitivity enhancer

The use of staining or contrast-enhanced media for enhancement of in vivo 3D visualization of tissues and cells architecture has been investigated with CT and MRI for angiography, mammography, and TE applications.^101,102 Due to its tunability (i.e., ability to select precise energies above and below the absorption edge of the contrast media of interest), contrast-enhanced KES may offer enhanced sensitivity for visualization and quantification of the complex 3D morphology and microarchitecture of soft and hard tissues. The effectiveness of KES has been evaluated for applications to mammography, coronary angiography, and cancer cell imaging.^52,57,59 However, the potential of KES (or KES-CT) for TE applications has not been considered. Labeling scaffolds or cells with bioinert contrast media with a K-edge energy in the range used in SR-biomedical imaging may improve the sensitivity for tracking cells, new tissue growth, and scaffold degradation kinetics. In some cases, contrast agents may also exert beneficial biochemical effects that may enhance cellular activity. For instance, both Sr- and barium-based materials have been studied as contrast media for KES. Interestingly, they have also been reported to influence cellular activity, material dissolution rate, bone remodeling, and provision of nonbridging oxygen that may optimize new tissue growth.^60,102,103 Unfortunately, Cooper et al.⁶⁰ noted that the relatively low K-edge of Sr (16.105 keV) may be a limitation, as this energy is lower than for most biomedical imaging methods. The use of GNPs as a contrast media to enhance SR-μCT has also been considered.^73,74 Astolfo et al.⁷⁴ report that the presence of GNPs facilitates a reduction in spatial resolution while maintaining the ability to localize the area of interest with acceptable image quality. Functionalized GNPs have also been reported to promote cell attachment, which is beneficial to TE applications.¹⁰³

Combining synchrotron techniques

To optimize the properties of constructs used in bone and cartilage TE, researchers often examine the functionality of their constructs in animal models. Even though longitudinal progressive monitoring of the functionality of the constructs is most desirable, they usually characterize the tissue ingrowth and properties of the implanted construct ex vivo. One of the prevalent techniques for examining excised samples is 2D SEM analysis.⁶⁷ SR techniques have different and interesting properties that may provide 3D cellular-level details in a manner comparable to SEM.⁶⁷ For example, inline PCI has a faster acquisition time, a simpler setup, and can offer nondestructive details of thin soft tissues.⁶⁶ DEI can effectively delineate thick samples with regions containing varying densities, including soft and hard tissues, and provide refraction, absorption, and extinction contrasts.¹⁷ Contrast-enhanced KES-CT or three-energy KES may improve sensitivity and enable characterization of construct degradation kinetics or tissue ingrowth. This may consequently enable effective delineation of the source of the newly growing tissues, that is, if they are from the host tissues or introduced labelled cells. Based on the capabilities of the different methods and the information obtainable, these methods can be combined for quantitative and qualitative examination of excised samples. Complementary use of DEI and inline PC-CT has been used to enable characterization of osteonecrosis repair.⁸⁴ Although using two SR-based imaging methods may provide more information about construct performance and the quality of integration of constructs with host tissues, this is only advisable for excised samples due to radiation dose and long exposure times posing risks to live animals and humans. However, the use of a source with a high flux and brightness, as discussed next, may serve to significantly reduce these concerns.¹²

Toward laboratory and clinical applications of synchrotron techniques in TE

Despite synchrotron light's interesting advantages, the scarcity of sources, their unavailability for routine clinical use, high costs, and associated complex instrumentation are challenges for researchers.⁴⁸ Researchers have engaged in the development of lab-based microfocus X-ray sources that have excellent sensitivity and increased contrast-to-noise ratio and which are suitable for clinical and biomedical applications.^{16,45–49,105} Wilkins and colleagues reported one of the earliest implementations of lab-based inline PCI²⁴, and successful implementations have since then been reported by many others.^105,106 Researchers have also developed DPCI with a conventional X-ray source and imaging detector.^45–47,107 Furthermore, research with regard to CAXPCI has focused primarily on the use of conventional X-ray sources and area detectors for laboratory or clinical PCI.^48–50

Though massive efforts have been directed toward the development of clinical PCI, to the best of the authors' knowledge, this progress is still best described as “moving towards,” and we look forward to commercialization of such systems.¹⁰⁷ Challenges that have limited the clinical implementation of PCI include beam-hardening artifacts and differential phase clipping.^106,109 Furthermore, the gantry of clinically used X-ray imaging systems cannot practically be lengthened beyond 1 meter; however, a longer source-to-detector distance is important for effective phase contrast.⁴⁹ The trade-off between image quality, exposure time, and absorbed dose is also a factor to be carefully considered for the clinical use of PCI methods.⁴⁸

Influence of synchrotron source

Flux and brilliance are the major markers of X-ray beam quality and vary from one synchrotron facility to another. The use of high-energy and high-brilliance sources generates a range of X-rays that comprise a hard and highly penetrating beam which is critical for biomedical imaging.^12,109 For progression to higher brightness, insertion devices such as a wiggler or undulator with the capability to optimize brilliance (number of photons/second/solid angle/bandwidth) by many orders of magnitude when compared with bending magnets may be used.¹² These increments in brightness mostly depend on the electron beam size, angular distribution, or excursion angle and may notably decrease the exposure time and dose used for experiments.¹² Hence, the use of insertion devices to boost brightness will facilitate the use of synchrotron sources for imaging live animals and possibly humans, optimize image quality, and reduce the exposure time and dose.^12,109

Conclusions

This review reports on technological developments, advantages, applications, and the potential of SR-based X-ray imaging techniques for bone and cartilage TE. SR-based imaging techniques are evolving as robust assessment techniques for TE as well as biomedical applications, in general, and could offer advantages over currently used evaluation techniques. For the high spatial resolution and quantitative imaging required for the characterization of engineered scaffolds and tissue growth ex vivo and in vitro, this review suggests exploring high-resolution SR-based imaging methods or combining different but complementary SR-based techniques. Using SR-based methods for the characterization of engineered scaffolds and tissue growth in vivo requires strategic adaptations or trade-offs; for example, enabling the acquisition of acceptable image quality at reduced radiation dose. In this regard, nanoparticles can be employed as contrast agents, the photon flux and brilliance can be increased, and the employed imaging technique can be modified, such as is done in AB-PCI compared with standard DEI. Although the use of SR-based imaging techniques is not yet as prevalent in TE as in biomedical imaging, representative examples of biomedical imaging applications of SR-based techniques are promising. The results from these works can be extrapolated to TE and used as a guide for tissue engineers with regard to choosing the right technique to suit their specific applications.

Footnotes

Acknowledgments

This work was supported by grants from the Saskatchewan Health Research Foundation (SHRF) and the Natural Sciences and Engineering Research Council of Canada (NSERC).

Disclosure Statement

The authors hereby declare that neither competing financial interests nor conflicts of interest exist.

References

, Othman

S.F.

, and Magin

R.L.

Monitoring tissue engineering using magnetic resonance imaging. J Biosci Bioeng, 106, 515, 2008.

Cancedda

, Cedola

, Giuliani

, Komlev

, Lagomarsino

, Mastrogiacomo

, et al. Bulk and interface investigations of scaffolds and tissue-engineered bones by X-ray microtomography and X-ray microdiffraction. Biomaterials, 28, 2505, 2007.

Marvel

, Okrasinski

, Bernacki

S.H.

, Loboa

, and Dayton

P.A.

Applications of low intensity pulsed ultrasound for functional bone tissue engineering using adult stem cells. Ultrasonics Symposium (IUS), 2009 IEEE International, 357, 2009.

Potter

, Sweet

D.E.

, Anderson

, Davis

G.R.

, Isogai

, Asamura

, et al. Non-destructive studies of tissue-engineered phalanges by magnetic resonance microscopy and X-ray microtomography. Bone, 38, 350, 2006.

, Gao

, Mao

, Zhou

, Shen

, Hu

, and Han

Collagen/chitosan porous scaffolds with improved biostability for skin tissue engineering. Biomaterials, 24, 4833, 2003.

Ismail

E.C.

, Kaabar

, Garrity

, Gundogdu

, Bunk

, Pfeiffer

, Farquharson

M.J.

, and Bradley

D.A.

X-ray phase contrast imaging of the bone–cartilage interface. Appl Radiat Isot, 68, 767, 2010.

Kowalski

, Kwan

H.H.

, Prionas

S.D.

, Allision

A.C.

, and Fajardo

L.F.

Characterization and applications of the disc angiogenesis system. Exp Mol Pathol, 56, 1, 1992.

Guldberg

R.E.

, Duvall

C.L.

, Peister

, Oest

M.E.

, Lin

A.S.P.

, Palmer

A.W.

, and Levenston

M.E.

3D imaging of tissue integration with porous biomaterials. Biomaterials, 29, 3757, 2008.

Khan

A.N.

Imaging of meniscus abnormalities in the temporomandibular joint: overview. Available at http://emedicine.medscape.com/article/385129-overview. (last accessed June 23, 2013).

10.

Zhu

, Chen

X.B.

, and Chapman

A brief review of visualization for nerve tissue engineering applications. Biomimetrics Biomater Tissue Eng, 7, 81, 2010.

11.

Hendee

W.R.

, Cleary

, Ehman

R.L.

, Fullerton

G.D.

, Grundfest

W.S.

, et al. Bioengineering and imaging research opportunities workshop V: summary of findings on imaging and characterizing structure and function in native and engineered tissues. Radiology, 248, 342, 2008.

12.

Chapman

L.D.

Health research using synchrotron techniques. Lecture on synchrotron radiation basics [PowerPoint slides]. Division of Biomedical Engineering University of Saskatchewan, course tools Blackboard. Available at https://bblearn.usask.ca, 2013 (last accessed June 23, 2013).

13.

Takeda

, Momose

, Wu

, Yu

, Zeniya

, Lwin

T.T.

, Yoneyama

, and Itai

Vessel imaging by interferometric phase-contrast x-ray technique. Circulation, 105, 1708, 2002.

14.

Chapman

, Thomlinson

, Johnston

R.E.

, Washburn

, Pisano

, Gmur

, et al. Diffraction enhanced X-ray imaging. Phys Med Biol, 42, 2015, 1997.

15.

Kunisada

, Shimao

, Sugiyama

, Takeda

, Ozaki

, and Ando

X-ray dark field imaging of human articular cartilage: possible clinical application to orthopedic surgery. Eur J Radiol, 68S, S18, 2008.

16.

Pfeiffer

, Weitkamp

, Bunk

, and David

Phase retrieval and differential phase-contrast imaging with low brilliance X-ray sources. Nat Phys, 2, 258, 2006.

17.

Izadifar

, Chapman

, and Chen

Computed tomography diffraction enhanced imaging for in situ visualization of tissue scaffolds implanted in cartilage. Tissue Eng Part C Methods, 20, 140, 2013.

18.

Zhu

, Chapman

, Cooper

, Schreyer

D.J.

, and Chen

X-ray diffraction enhanced imaging as a novel method to visualize low-density scaffolds in soft tissue engineering. Tissue Eng Part C Methods, 17, 1071, 2011.

19.

Gao

, Pogany

, Stevenson

A.W.

, and Wilkins

S.W.

Phase-contrast radiography. Radiographics, 18, 1257, 1998.

20.

Greaves

A.N.

Phase Contrast in X-rays Imaging [Doctoral thesis]. Swinburne University of Technology, Melbourne, Australia. 2010.

21.

Zernike

Nobel lecture: how I discovered phase contrast [Model Lecture]. Science, New Series, 121, 345, 1953. (last accessed June 24, 2013).

22.

Appel

, Anastasio

M.A.

, and Brey

E.M.

Potential for imaging engineered tissues with X-ray phase contrast. Tissue Eng Part B Rev, 17, 321, 2011.

23.

Gabor

A new microscopic principle. Nature, 161, 777, 1948.

24.

Wilkins

S.W.

, Gureyev

T.E.

, Gao

, Pogany

, and Stevenson

A.W.

Phase-contrast imaging using polychromatic hard X-rays. Nature, 384, 335, 1996.

25.

Spanne

, Raven

, Snigireva

, and Snigirev

In-line holography and phase-contrast microtomography with high energy x-rays. Phys Med Biol, 44, 741, 1999.

26.

Krol

In-line hard X-ray holography for biomedical imaging. In: Rosen

., ed. Holography, Research and Technologies. 2011. Available at http://cdn.intechopen.com/pdfs/13863/InTechIn_line_hard_x_ray_holography_for_biomedical_imaging.pdf (last accessed June 24, 2013).

27.

Kieffer

J.C.

, Krol

, Jiang

, Chamberlain

C.C.

, Scalzetti

, and Ichalalene

Future of laser-based X-ray sources for medical imaging. J Appl Phys B, 74, S75, 2002.

28.

Zhou

S.A.

, and Brahme

Development of phase-contrast x-ray imaging techniques and potential medical applications. Phys Med, 24, 129, 2008.

29.

Bonse

, and Hart

An X-ray interferometer with Bragg case beam splitting and beam recombination. Z Phys, 194, 155, 1966.

30.

Yoneyama

, Yamada

, and Takeda

Fine biomedical imaging using X-ray phase-sensitive technique. In: Gargiulo

G.D.

, and McEwan

, eds. Advanced Biomedical Engineering. 2011. Available at www.intechopen.com/books/authors/advanced-biomedical-engineering/fine-biomedical-imaging-using-x-ray-phase-sensitive-technique (last accessed June 25, 2013).

31.

Grubel

, Stephenson

G.B.

, Gutt

, Sinn

, and Tschentscher

XPCS at the European X-ray free electron laser facility. Nucl Instrum Methods Phys Res, B262, 357, 2007.

32.

Momose

Phase-contrast X-ray imaging based on interferometry. J Synchrotron Radiat, 9, 136, 2002.

33.

Ingal

V.N.

, and Beliaevskaya

E.A.

X-ray plane-wave topography observation of the phase contrast from a non-crystalline object. J Phys D Appl Phys, 28, 2314, 1995.

34.

Davis

T.J.

, Gao

, Gureyev

T.E.

, Stevenson

A.W.

, and Wilkins

S.W.

Phase-contrast imaging of weakly absorbing materials using hard X-rays. Nature, 373, 595, 1995.

35.

Zhong

, Thomlinson

, Chapman

, and Sayers

Implementation of diffraction-enhanced imaging experiments at the NSLS and APS. Nucl Instrum Methods Phys Res, 450, 556, 2000.

36.

Hasnah

M.O.

, Zhong

, Oltulu

, Pisano

, Johanston

R.E.

, Sayers

, et al. Diffraction enhanced imaging contrast mechanisms in breast cancer specimens. Med Phys, 29, 2216, 2002.

37.

Rigon

, Arfelli

, and Menk

R.H.

Three-image diffraction enhanced imaging algorithm to extract absorption, refraction, and ultrasmall-angle scattering. Appl Phys Lett, 90, 14102, 2007.

38.

Chou

C.Y.

, Anastasio

M.A.

, Brankov

J.G.

, Wernick

M.N.

, Brey

E.M.

, Connor

D.M.

, and Zhong

An extended diffraction-enhanced imaging method for implementing multiple-image radiography. Phys Med Biol, 52, 1923, 2007.

39.

Khelashvili

, Brankov

J.G.

, Chapman

, Anastasio

M.A.

, Yang

, Zhong

, and Wernick

M.N.

A physical model of multiple-image radiography. Phys Med Biol, 51, 21, 2006.

40.

, Zhong

, Connor

, Mollenhauer

, and Muehleman

Phase-sensitive X-ray imaging of synovial joints. Osteoarthritis Cartilage, 17, 1193, 2009.

41.

Sunaguchi

, Yuasa

, Huo

, Ichihara

, and Ando

X-ray refraction-contrast computed tomography images using dark-field imaging optics. Appl Phys Lett, 97, 153701, 2010.

42.

Revol

, Plank

, Kaufmann

, Kastner

, Kottler

, and Neels

Laminate fibre structure characterisation of carbon fibre-reinforced polymers by X-ray scatter dark-field imaging with a grating interferometer. NDT&E Int, 58, 64, 2013.

43.

Pfeiffer

, Bech

, Bunk

, Kraft

, Eikenberry

E.F.

, Brönnimann

, Grünzweig

, and David

Hard X-ray dark-field imaging using a grating interferometer. Nat Mater, 7, 134, 2008.

44.

, Zambelli

, Bevins

, Ge

, and Chen

G.H.

Spatial resolution characterization of differential phase contrast CT systems via modulation transfer function (MTF) measurements. Phys Med Biol, 58, 4119, 2013.

45.

Stutman

, Beck

T.J.

, Carrino

J.A.

, and Bingham

C.O.

Talbot phase-contrast x-ray imaging for the small joints of the hand. Phys Med Biol, 56, 5697, 2011.

46.

Chen

G.H.

, Bevins

, Zambelli

, and Qi

Small-angle scattering computed tomography (SAS-CT) using a Talbot-Lau interferometer and a rotating anode x-ray tube: theory and experiments. Opt Express, 18, 12960, 2010.

47.

Hauser

, Wang

, Kubik-Huch

R.A.

, Trippel

, Singer

, Hohl

M.K.

, et al. A study on mastectomy samples to evaluate breast imaging quality and potential clinical relevance of differential phase contrast mammography. Invest Radiol, 49, 131, 2014.

48.

Olivo

, and Speller

R.A.

Coded-aperture technique allowing x-ray phase contrast imaging with conventional sources. Appl Phys Lett, 91, 074, 2007.

49.

Munro

P.R.T.

, Ignatyev

, Speller

R.D.

, and Olivo

Source size and temporal coherence requirements of coded aperture type x-ray phase contrast imaging systems. Opt Express, 18, 19681, 2010.

50.

Krejci

, Jakubek

, and Kroupa

X-ray phase contrast imaging using single absorption coded aperture. Nucl Instrum Methods Phys Res, A633, S181, 2011.

51.

Kavanagh

, and Vojnovic

Phase contrast x-ray imaging. 2011. Available at http://users.ox.ac.uk/∼atdgroup/technicalnotes/Phase%20Constrast%20X-ray%20imaging.pdf (last accessed 25/06/2013).

52.

Bewer

, Zhang

, Zhu

, Zhang

, George

G.N.

, Pickering

I.J.

, and Chapman

Development of a combined K-edge subtraction and fluorescence subtraction imaging system for small animals. Rev Sci Instrum, 79, 085102, 2008.

53.

Jacobson

Dichromatic absorption radiography; dichromography. Acta Radiol, 39, 437, 1953.

54.

Sarnelli

, Taibi

, Tuffanelli

, Baldazzi

, Bollini

, Cabal Rodriguez

A.E.

, Gombia

, Prino

, Ramello

, Tomassi

, and Gambaccini

K-edge digital subtraction imaging based on a dichromatic and compact x-ray source. Phys Med Biol, 49, 3291, 2004.

55.

Zhong

, Chapman

, Menk

, Richardson

, Theophanis

, and Thomlinson

Monochromatic energy subtraction radiography using a rotating anode source and a bent Laue monochromator. Phys Med Biol, 42, 1751, 1997.

56.

Zhu

Bent Laue Crystals in Biomedical X-ray Imaging Applications [Doctoral dissertation]. University of Saskatchewan, Saskatoon, Canada, 2012. Available at http://ecommons.usask.ca/handle/10388/ETD-2012-11-768 (last accessed June 26, 2013).

57.

Bayat

, Le Duc

, Porra

, Berruyer

, Nemoz

, Monfraix

, et al. Quantitative functional lung imaging with synchrotron radiation using inhaled xenon as contrast agent. Phys Med Biol, 46, 3287, 2001.

58.

Riederer

S.J.

, Kruger

R.A.

, and Mistretta

C.A.

Three-beam K-edge imaging of iodine using differences between fluoroscopic video images: theoretical considerations. Med Phys, 8, 471, 1981.

59.

Lewin

J.M.

, Isaacs

P.K.

, Vance

, and Larke

F.J.

Dual-energy contrast enhanced digital subtraction mammography: feasibility. Radiology, 229, 261, 2003.

60.

Cooper

D.M.

, Chapman

L.D.

, Carter

, Wu

, Panahifar

, Britz

H.M.

, et al. Three-dimensional mapping of strontium in bone by dual energy K-edge subtraction imaging. Phys Med Biol, 57, 5777, 2012.

61.

Feldkamp

L.A.

, Goldstein

S.A.

, Parfitt

A.M.

, Jesion

, and Kleerekoper

The direct examination of three-dimensional bone architecture in vitro by computed tomography. J Bone Miner Res, 4, 3, 1989.

62.

Cooper

D.M.

, Turinsky

, Sensen

, and Hallgrimsson

Effect of voxel size on 3D micro-CT analysis of cortical bone porosity. Calcif Tissue Int, 80, 211, 2007.

63.

Peter

Z.A.

, and Peyrin

Synchrotron radiation micro-CT imaging of bone tissue. In: Homma

, ed. Theory and Applications of CT Imaging and Analysis. 2011. Available at www.intechopen.com/books/theory-and-applications-of-ct-imaging-and-analysis/synchrotron-radiation-micro-ct-imaging-of-bone-tissue (last accessed July 10, 2013).

64.

Stock

S.R.

Recent advances in X-ray microtomography applied to materials. Int Mat Rev, 53, 129, 2008.

65.

Cooper

D.M.L.

, Erickson

, Peele

A.G.

, Hannah

, Thomas

C.D.L.

, and Clement

J.G.

Visualization of 3D osteon morphology by synchrotron radiation micro-CT. J Anat, 219, 481, 2011.

66.

Yue

, Lee

P.D.

, Poologasundarampillai

, Yao

, Rockett

, Devlin

A.H.

, et al. Synchrotron X-ray microtomography for assessment of bone tissue scaffolds. J Mater Sci Mater Med, 21, 847, 2010.

67.

Peyrin

, Mastrogiacomo

, Cancedda

, and Martinetti

SEM and 3D Synchrotron radiation micro-tomography in the study of bioceramic scaffolds for tissue-engineering applications. Biotechnol Bioeng, 97, 638, 2007.

68.

Albertini

, Giuliani

, Komlev

, Moroncini

, Pugnaloni

, Pennesi

, et al. Organization of extracellular matrix fibers within polyglycolic acid-polylactic acid scaffolds analyzed using X-ray synchrotron-radiation phase-contrast micro computed tomography. Tissue Eng Part C Methods, 15, 403, 2009.

69.

Zehbe

, Haibel

, Schmidt

, Riesemeier

, Kirkpatrick

C.J.

, Schubert

, and Brochhausen

High-resolution x-ray tomography—3D imaging for tissue engineering applications. In: Eberli

, ed. Tissue Engineering. 2011. Available at http://cdn.intechopen.com/pdfs/9794/InTech-High_resolution_x_ray_tomography_3d_imaging_for_tissue_engineering_applications.pdf (last accessed July 10, 2013).

70.

Mastrogiacomo

, Komlev

V.S.

, Hausard

, Peyrin

, Turquier

, Casari

, et al. Synchrotron radiation microtomography of bone engineered from bone marrow stromal cells. Tissue Eng, 10, 1767, 2004.

71.

Papadimitropoulos

, Mastrogiacomo

, Peyrin

, Molinari

, Komlev

V.S.

, Rustichelli

, and Cancedda

Kinetics of in vivo bone deposition by bone marrow stromal cells within a resorbable porous calcium phosphate scaffolds: a X-ray computed microtomography study. Biotechnol Bioeng, 98, 271, 2007.

72.

Komlev

V.S.

, Peyrin

, Mastrogiacomo

, Cedola

, Papadimitropoulos

, Rustichelli

, and Cancedda

Kinetics of in vivo bone deposition by bone marrow stromal cells into porous calcium phosphate scaffolds: an X-ray computed microtomography study. Tissue Eng, 12, 3449, 2006.

73.

Astolfo

, Arfelli

, Schultke

, James

, Mancinie

, and Menke

R.H.

A detailed study of gold-nanoparticle loaded cells using X-ray based techniques for cell-tracking applications with single-cell sensitivity. Nanoscale, 5, 3337, 2013.

74.

Astolfo

, Schültke

, Menk

R.H.

, Kirch

R.D.

, Juurlink

B.H.J.

, Hall

, et al. In vivo visualization of gold-loaded cells in mice using x-ray computed tomography. Nanomedicine, 9, 284, 2013.

75.

Weiss

, Obadia

, Magne

, Bourges

, Rau

, Weitkamp

, et al. Synchrotron X-ray microtomography (on a micron scale) provides three-dimensional imaging representation of bone ingrowth in calcium phosphate biomaterials. Biomaterials, 24, 4591, 2003.

76.

Komlev

V.S.

, Mastrogiacomo

, Peyrin

, Cancedda

, and Rustichelli

X-ray synchrotron radiation pseudo-holotomography as a new imaging technique to investigate angio- and microvasculogenesis with no usage of contrast agents. Tissue Eng Part C Methods, 15, 425, 2009.

77.

Issever

A.S.

, Diederichs

, Majumdar

, Rogalla

, Hamm

B.K.

, Lange

, et al. Analyser-based tomography images of cartilage. J Synchrotron Radiat, 15, 525, 2008.

78.

Muehleman

, Majumdar

, Issever

, Arfelli

, Menk

R.H.

, Rigon

, et al. X-ray detection of structural orientation in human articular cartilage. Osteoarthritis Cartilage, 12, 97, 2004.

79.

Wagner

, Aurich

, Sieber

, Stoessel

, Wetzel

W.D.

, Schmuck

, et al. Options and limitations of joint cartilage imaging: DEI in comparison to MRI and sonography. Nucl Inst Methods Phys Res A 548, 47, 2005.

80.

Muehleman

, Fogarty

, Reinhart

, Tzvetkov

, Li

, and Nesch

In-laboratory diffraction-enhanced X-ray imaging for articular cartilage. Clin Anat, 23, 530, 2010.

81.

Mollenhauer

, Aurich

M.E.

, Zhong

, Muehleman

, Cole

A.A.

, Hasnah

, Oltulu

, Kuettner

K.E.

, Margulis

, and Chapman

L.D.

Diffraction-enhanced X-ray imaging of articular cartilage. Osteoarthritis Cartilage, 10, 163, 2002.

82.

Ando

, Hashimoto

, Hashizume

, Hyodo

, Inoue

, Kunisada

, et al. Clinical step onward with X-ray dark-field imaging and perspective view of medical applications of synchrotron radiation in Japan. Nucl Instrum Methods Phys Res A 548, 1, 2005.

83.

Bech

, Jensen

T.H.

, Bunk

, Donath

, David

, Weitkamp

, et al. Advanced contrast modalities for X-ray radiology: phase-contrast and dark-field imaging using a grating interferometer. Z Med Phys, 20, 7, 2010.

84.

Sun

, Li

Z.R.

, Yang

Y.R.

, Shi

Z.C.

, Wang

, Liu

, and Shi

Experimental study on phase-contrast imaging with synchrotron hard X-ray for repairing osteonecrosis of the femoral head. Orthopedics, 34, e530, 2011.

85.

Zehbe

, Haibel

, Riesemeier

, Gross

, Kirkpatrick

C.J.

, Schubert

, and Brochhausen

Going beyond histology: synchrotron μCT as a complementary methodology for biological tissue characterization. J R Soc Interface, 7, 49, 2010.

86.

Thurner

, Karamuk

, and Muller

3D characterization of fibroblast cultures on PET-textiles. Eur Cell Mater, 2, 57, 2001.

87.

Muller

, Riedel

, and Thurner

P.J.

Three-dimensional characterization of cell clusters using synchrotron-radiation-based micro-computed tomography. Microsc Microanal, 12, 97, 2006.

88.

Zehbe

, Geobbels

, Ibold

, Gross

, and Schubert

Three-dimensional visualization of in vitro cultivated chondrocytes inside porous gelatine scaffolds: a tomographic approach. Acta Biomater, 6, 2097, 2010.

89.

Choi

C.H.

, Kim

H.T.

, Choe

J.Y.

, Kim

S.K.

, Choi

G.W.

, Jheon

, and Kim

J.K.

In vivo high-resolution synchrotron radiation imaging of collagen-induced arthritis in a rodent model. J Synchrotron Radiat, 17, 393, 2010.

90.

Muehleman

, Li

, and Zhong

Preliminary study on diffraction enhanced radiographic imaging for a canine model of cartilage damage. Osteoarthritis Cartilage, 14, 882, 2006.

91.

Coan

, Bamberg

, Diemoz

P.C.

, Bravin

, Timpert

, Mützel

, et al. Characterization of osteoarthritic and normal human patella cartilage by computed tomography X-ray phase-contrast imaging: a feasibility study. Invest Radiol, 45, 437, 2010.

92.

Coan

, Wagner

, Bravin

, Diemoz

P.C.

, Keyrilainen

, and Mollenhauer

In vivo x-ray phase contrast analyzer-based imaging for longitudinal osteoarthritis studies in guinea pigs. Phys Med Biol, 55, 7649, 2010.

93.

Cooper

D.M.

, Bewer

, Wiebe

, Wysokinski

T.W.

, and Chapman

Diffraction enhanced X-ray imaging of the distal radius: a novel approach for visualization of trabecular bone architecture. Can Assoc Radiol J, 62, 255, 2011.

94.

Muehleman

, Chapman

L.D.

, Kuettner

K.E.

, Rieff

, Mollenhauer

J.A.

, Massuda

, and Zhong

Radiography of rabbit articular cartilage with diffraction enhanced imaging. Anat Rec, 272, 392, 2003.

95.

Connor

D.M.

, Sayers

, Sumner

D.R.

, and Zhong

Diffraction enhanced imaging of controlled defects within bone, including bone-metal gaps. Phys Med Biol, 51, 3283, 2006.

96.

Connor

D.M.

, Hallen

H.D.

, Lalush

D.S.

, Sumner

D.R.

, and Zhong

Comparison of diffraction-enhanced computed tomography and monochromatic synchrotron radiation computed tomography of human trabecular bone. Phys Med Biol, 21, 6123, 2009.

97.

Thurner

, Muller

, Raeber

, Sennhauser

, and Hubbel

J.A.

3D morphology of cell cultures: a quantitative approach using micrometer synchrotron light tomography. Microsc Res Tech, 66, 289, 2005.

98.

Zehbe

, Haibel

, Brochhausen

, Gross

, Kirkpatrick

C.J.

, and Schubert

Characterization of oriented protein-ceramic and protein-polymer-composites for cartilage tissue engineering using synchrotron m-CT. Int J Mater Res, 98, 562, 2007.

99.

, Zhong

, Lidtke

, Kuettner

K.E.

, Peterfy

, Aliyeva

, and Muehleman

Radiography of soft tissue of the foot and ankle with diffraction enhanced imaging. J Am Podiatr Med Assoc, 94, 315, 2004.

100.

Bolland

B.J.R.F.

, Kanczler

J.M.

, Ginty

P.J.

, Howdle

S.M.

, Shakesheff

K.M.

, Dunlop

D.G.

, and Oreffo

R.O.C.

The application of human bone marrow stromal cells and poly (DL-lactic acid) as a biological bone graft extender in impaction bone grafting. Biomaterials, 29, 3221, 2008.

101.

Lao

, Nedelec

J.M.

, and Jallot

New strontium based bioactive glasses: physicochemical reactivity and delivering capability of biologically active dissolution products. J Mater Chem, 19, 2940, 2009.

102.

Kaur

, Sharma

, Kumar

, and Singh

Assessment of in-vitro bioactivity of SiO2-BaOZnO-B2O3-Al2O3 glasses: an optico-analytical approach. Mat Sci Eng Part C, 32, 1941, 2012.

103.

Jung

, Minami

, Patel

, Lee

, Jiang

, Yuan

, et al. Incorporation of functionalized gold nanoparticles into nanofibers for enhanced attachment and differentiation of mammalian cells. J Nanobiotechnol, 11, 10, 2012.

104.

Kashyap

Y.S.

, Yadav

P.S.

, Roy

, Sarkar

P.S.

, Shukla

, and Sinha

Laboratory-based X-ray phase-contrast imaging technique for material and medical science applications. Appl Radiat Isot, 66, 1083, 2008.

105.

Kelly

K.L.

, and Tanner

B.K.

Factors affecting in-line phase contrast imaging with a laboratory microfocus x-ray source. 2006. Available at www.icdd.com/resources/axa/vol49/V49_04.pdf (last accessed July 10, 2013).

106.

Stampanoni

, Wang

Z.T.

, Thuring

, David

, Rossel

, Trippel

, et al. The first analysis and clinical evaluation of native breast tissue using differential phase-contrast mammography. Invest Radiol, 46, 801, 2011.

107.

Bravin

, Coan

, and Suortti

X-ray phase-contrast imaging: from pre-clinical applications towards clinics. Phys Med Biol, 58, R1, 2013.

108.

Tanaka

, Nagashima

, Kido

, Hoshino

, Kiyohara

, Makifuchi

, Nishino

, Nagatsuka

, and Momose

Cadaveric and in vivo human joint imaging based on differential phase contrast by X-ray Talbot-Lau interferometry. Z Med Phys, 20, 222, 2013.

109.

Argonne National Laboratory. The advanced photon source (APS) overview: Insertion devices and brilliance. Available at www.aps.anl.gov/About/APS_Overview/insertion_devices.html (last accessed July 10, 2013).