Magnetic Resonance Imaging of Hard Tissues and Hard Tissue Engineered Bio-substitutes

Magnetic resonance imaging (MRI) is a non-invasive imaging technology that performs three-dimensional (3D) imaging for body organs with high spatial resolution (< 1 mm). The primary advantage of MRI is to allow for obtaining cross-sectional anatomical information of the organs due to an excellent soft tissue contrast and excellent depth penetration. Furthermore, MRI offers functional information based on various tissue properties, such as proton density, temperature, biochemical content, oxygen level, and pH [1, 2].

Since MRI has been applied in the clinic in early 1980s, the number of MR examinations has tremendously increased especially for pathologies related to brain, spine, abdomen, cardiovascular, and muscular systems [3, 4]. To date, the state-of-art MRI techniques include 1) MR spectroscopy (MRS) and chemical shift imaging (CSI), which permit the identification and classification of tumors through the assessment of the chemical metabolism in a specific area of the body; 2) functional MRI (fMRI), which is able to image the neuronal activity through the detection of the blood oxygenation level-dependent (BOLD) changes occurring in the brain; and 3) MR elastography (MRE), which can be used for the characterization of the biomechanical properties of soft tissues by the visualization of propagating shear waves [5‐7].

Despite the wide use of MRI as a diagnostic tool for soft tissues, its application for hard tissues (i.e., bone and teeth) remains challenging. The water content in such tissues is usually very low (< 20 % v/v) and mainly trapped in a solid phase leading to a very short relaxation profile. Therefore, the signal decays significantly before the detection by conventional MRI techniques, which results in a dark and undefined image [8]. Over the last decade, newly developed MRI techniques that are able to image tissues with very low water content and ultrashort T2 relaxation times have been reported. The most promising techniques include ultrashort echo time (UTE) imaging, zero echo time (ZTE) imaging, and sweep imaging with Fourier transformation (SWIFT).

Thus, after a general introduction about basic principles of MRI, bone-specific MRI sequences, and anatomical and chemical compositions of hard tissues, this review mainly emphasizes all the progresses that have been made on ultrashort MRI acquisitions of such tissues in terms of sequence development, possible clinical translations, and related strengths and weaknesses. Finally, this review focused also on a fast-growing application of such MRI sequences, i.e., acquisition and in vivo monitoring of biomaterials intended to bone and dental restoration.

MR imaging was evolved from the basic physical principles of nuclear magnetic resonance (NMR). Typically, MRI is used to detect the nucleus of hydrogen (H⁺) in the water molecules as present in the body. Protons are positively charged subatomic particles that are constantly rotating along their axis. The rotational movement of protons in an atom gives rise to a quantum-mechanic atomic property that is called spin. Only unpaired protons exhibit spin, which is the case of atoms with an odd number of protons and neutrons, whereas atoms with an equal number of protons and neutrons exhibit a null overall spin. The spinning motion of an electric charge generates a magnetic field, hence each proton can be considered as a small bar magnet and described as a vectoral force, named magnetic moment, with a random orientation (Fig. 1a). When an external magnetic field (B₀) is applied to a proton, its magnetic moment will align to the direction of B₀ and have only two possible orientations, i.e., parallel to B₀ (spin = 1/2), and antiparallel to B₀ (i.e., spin = −1/2). The speed of the precession movement of the proton is described by the Larmor frequency (ω) with unit of rad/s, and depends on the strength of the applied B₀ with unit of Tesla (T), and on gyromagnetic ratio (γ) according to the following formula ω = γ B₀, where γ is a constant describing the gyromagnetic ratio of a particular nuclear species (i.e., for proton γ = 267.56 × 10⁶ rad/s/T). If to the static B₀ a second alternating magnetic field (B₁) is applied in the orthogonal direction, it is possible to disturb the magnetic moment of the proton from the B₀ direction. The alternating magnetic field B₁ is usually referred as radio frequency (RF) and must have the same Larmor frequency as the investigated nucleus. In this way, the applied RF is able to induce a transfer of energy to the protons with a rotation angle determined by the amplitude and duration of B1 (Fig. 1a). When the RF is switched off, the protons tend to return to their equilibrium status in a process known as relaxation.

During relaxation, protons lose energy according to two processes known as longitudinal relaxation (T₁) and transverse relaxation (T₂) (Fig. 1b, c). The longitudinal relaxation describes the process whereby the adsorbed energy is released to the surrounding nuclei, inducing an increase of vibration within the lattice that results in an increase of thermal energy. Hence, longitudinal relaxation is also named spin-lattice relaxation, while T₁ relaxation refers to the time needed by the magnetization moment M_z to reach 63 % of the original position (i.e., M_0, Fig. 1b, d). Transverse relaxation describes the process whereby the adsorbed energy is released by spins as a consequence of random mutual interference between each other, hence also known as spin-spin relaxation. T₂ relaxation refers to the time needed by the magnetization moment M_xy to decay to a 37 % of its initial value (Fig. 1c, e). Pure T₂ decay is happening theoretically only with a completely homogeneous B₀. However, this is never the case, as tissues with different susceptibility properties can affect the homogeneity of the magnetic field. These inhomogeneities are described by a T₂^* constant (Fig. 1e). As tissues have different relaxation properties, an image can be generated based on the differences in T₁ (i.e., T₁-weighted image, Fig. 1d) and in T₂ (i.e., T₂-weighted image, Fig. 1e) relaxation properties [1, 9, 10].

The position of the spins into the space is defined through three separate magnetic field gradients named slice-selection gradient, phase-encoding gradient (G_P) and frequency encoding gradient (G_F). Firstly, a slice into the object is selected through the slice-selection gradient that generates a gradient field along the chosen axis, thus altering B₀ in the chosen direction. Usually, a RF is applied to excite a slice with a defined thickness which makes the protons precess in phase. Afterwards, the G_P is applied, which induce the protons to precess at different speeds according to their positions. When G_P is turned off, the protons have the same precession frequency, but they will be no longer in phase. Finally, G_F is applied perpendicularly along G_P direction inducing the protons to rotate at different frequencies according to their relative positions. A mathematical algorithm known as Fourier transform (FT) is used to define the position of the protons based on the frequency analysis of the measured MRI signal [9]. Specifically, FT is able to convert the frequency data (in Hz) in spatial information (in mm) that are collected in the so-called k-space. In a 2D acquisition, the k-space consists of a two-dimensional matrix with Kx and Ky coordinates corresponding to the x- and y-axis, respectively. Therefore, the location of the spins can be defined based on their spatial frequency in the Kx and Ky directions of the k-space [11, 12].

Tissues with a high free water content, such as blood, brain, skeletal muscle, and spinal cord, show long T₂ and T₁ relaxation times, which can generate a relatively strong signal in MRI [13]. By contrast, in tissues with low water content where the protons are trapped in a rigid crystal phase, such as bone and teeth, the signal decays quickly, resulting to a dark image with no discernible structured contrast (Table 1) [8, 14]. This review focused on the tissues with short T₂ relaxation time that are commonly classified as tissue with short (i.e., T₂ = 1–10 ms), ultrashort (i.e., T₂ = 0.1–1 ms), and supershort relaxation (i.e., T₂ < 0.1 ms) [15].

Hard tissues comprise all the tissues that show a mineralized component in their extracellular matrix, such as bone and teeth. The ratio of the mineral phase is not only different for each tissue but can vary with age, sex, gender, and site [50]. Such tissue heterogeneity results in different T₂ properties according to the amount of available free water (Table 1). In the following paragraphs, a general overview of the specific chemical components of bone and teeth is provided.

Conventional UTE (CUTE) as well as optimized UTE sequences were firstly used to image human osseous tissues (e.g., periosteum, knee, lumbar spine, cortical bone in the tibia) by Bydder’s group (Table 2) [28, 29]. The implementations of the CUTE sequence were based on two different strategies: reduction of the signal from long-T₂ components and fat suppression. Through the reduction of the signal from long-T₂ components an increase of the image contrast of the short-T₂ components was obtained, while fat suppression allowed a reduction of the background signal. Based on these strategies different optimized UTE sequences were developed, such as long-T₂ suppression UTE (LUTE), fat suppressed UTE (FUTE), short inversion time (TI) with UTE (STUTE), and medium TI with UTE (MUTE, Table 2). Among them, FUTE was found to be the most promising sequence as it was able to show contrast of the tibial periosteum and of the cortex in healthy patients. However, the signal from the cortical bone, which was associated to the collagen type I and the bound water, was found to decrease with the increase of the distance of the tissue from the MRI coil [29].

A different approach to allow UTE imaging in clinical applications was based on the subtraction of subsequent UTE images acquired with longer echo times. The subtracted UTE images resulted in a reduced signal from the tissue with long T₂ and highlighted the signal contrast from the short T₂ tissues. In this way it was possible to obtain high signal to noise images of the periosteum and of the bone at different locations in human volunteer (e.g., tibia, fibula, and spine). Interestingly, the thickness of the periosteum obtained from MRI was reported being consistent with the standard anatomical descriptions [69]. Furthermore, this subtraction method was used to obtain a positive signal from the cortical bone with higher conspicuity. Specifically, subtracted images acquired with FUTE sequence was used to show callus formation in a fractured tibia of a 22-year-old male patient 3 weeks after the injury, as well as the difference in cortical bone density between healthy and osteoporotic patients. On the other hand, subtracted images acquired with CUTE showed new bone formation occurring in 32-year-old female patients after 4 years from a tibia injury [70]. Nevertheless, these modified UTE sequences were subject to susceptibility and gradient distortion artifacts. Suppression of the signal from long-T₂ and fat components was also evaluated by combining the UTE sequence with an off-resonance saturation contrast (UTE-OSC, Table 2). The long T₂ components are minimally affected by the off-resonance saturation pulse; hence, the subtraction of UTE images with and without OSC resulted in water and fat suppression that selectively depicted the short T₂ components [30].

UTE spectroscopic imaging methods (UTESI) were developed based on highly undersampled interleaved projection reconstructions with multi-echo time sequences (Table 2) [71, 72]. Using the UTESI sequence, it was possible to achieve high spatial resolution spectroscopic images of the cortical bone in the order of 0.08–0.27 mm³, which is higher than the conventional spectroscopic techniques (in the order of mm³). The success of this spectroscopic technique arises from three main factors: 1) a minimal TE of 8 μs that was achieved through the combination of a variable-rate selective excitation (VERSE), a radial ramp selection, and a fast transmit/receive switch; 2) an interleaved variable TE UTE acquisition that significantly improved the spatial resolution; 3) suppression of long T₂ components from muscle and fat.

Further improvements of UTE imaging methods consisted in a dual adiabatic inversion recovery UTE sequence (DIR-UTE) and an adiabatic inversion recovery UTE sequence (IR-UTE), which were tested for cortical bone (Table 2). DIR-UTE method was able to suppress long-T₂ component and the signal from the surrounding fat and muscle, displaying the distal tibia of healthy volunteer with high contrast. However, DIR-UTE is a 2D technique, hence subject to partial-volume effect when tiny structures (< 0.1 mm) need to be images [73]. IR-UTE method provided excellent qualitative and quantitative depiction of the bone in vitro and in vivo, based on a robust long T₂ species suppression achieved using an inversion recovery and a short TR (Table 2). In comparison with a standard UTE sequence, IR-UTE technique provided lower SNR but higher CNR [31].

Importantly, the possibility to quantify the proton density in hard tissues through UTE-based approaches opened a new scenario for MRI. In fact, UTE- and ZTE-based MRI could be used as a screening tool for the detection of bone diseases related to bone pore size and water content changes (e.g., in osteoporotic conditions), as well as for understanding the healthy condition of the bone. It is known that a deficit of bone mineral content is related to an increase in water content; an increase in free bulk porous water corresponds to increased bone porosity; tightly-bound water is associated with the mineral content [74]. Gravimetric studies, ¹H NMR, and three-point bending test performed on tibiae harvested from healthy and phosphorous-deficient rabbits, have proven that bound water is associated with bone strength and toughness, while free water is associated with the elasticity modulus of bone. Therefore, mineral features and water content can be related to the bone biomechanical properties, and to the risk of bone fractures [74].

Chang et al. combined the UTE sequence with the magnetization transfer (MT), which is an indirect imaging method that allows a quantitative and qualitative assessment of protons with extremely fast relaxation [32]. UTE-MT imaging provided accurate information about cortical bone tissue properties by using a whole-body clinical 3 T MRI scanner in only 2 min acquisition time. The off-resonance saturation ratio (OSR) measured through UTE-MT method on human cadaveric femora and tibia correlates with the bone porosity measured via μCT and the mechanical properties, such as Young’s modulus and yield stress, measured through a four-point bending test. Specifically, a moderate negative correlation between the OSR and the cortical porosity (R² = 0.51), and a positive correlation between OSR and both Young’s modulus (R² = 0.12) and yield stress (R² = 0.30), were observed [32]. Horch et al. associated a preexisting UTE pulses with a double adiabatic full passage (DAFP) sequence and an adiabatic inversion recovery (AIR) for the detection of the pore and bound water, respectively. These sequences were used to acquire human cadaveric femora on a 4.7 T system. The estimated water content was correlated to the mechanical properties, such as Young’s modulus, yield stress, peak stress, and toughness to failure measured through three-point bending tests. Pearson’s correlation values between DAFP- and AIR-UTE method and the various mechanical properties were ranging between 0.35 and 0.69, hence providing a proof-of-concept that these UTE-based strategies can be used for the prediction of bone fractures [75]. Afterwards, the same authors translated the DAFP and AIR UTE protocols for the quantitative mapping of bound and porous water in radius and tibia of six healthy volunteers on a 3 T MRI clinical system. The outcomes of the study showed that the bound water concentration was approximately 28 and 35 mol ¹H/l of bone, while the porous water was 7 and 6 mol ¹H/l of bone, for tibia and radius respectively. These findings corroborated with the values reported in previous observations in femoral specimens [76]. Recently, a few more UTE-based methods to detect water content in vivo with more accuracy and precision have been proposed, such as IR-UTE, 3D IR-UTE, and UTE sequence associated with a numerical-based algorithm [77‐79]. However, the statistical correlation of these methodologies with the mechanical properties remains unexplored. Furthermore, water molecules bound to the collagen matrix or to the crystal phase are decaying really fast (< 10 μs) and their detection is still challenging even by using the most advanced UTE sequence.

The feasibility to use 3D UTE-MRI for the depiction of the mandibular condyle morphology, which is indicative of degenerative joint diseases, was investigated on human cadaveric specimens and compared with measurements performed via μCT. UTE-MRI was able to show the contour of the condyle cortical bone and the associated cartilage, which was not visible in μCT images. MRI reported accurate information, with an isotropic voxel size of 100 μm, including joint surface curvature, incongruity, as well as cartilage thickness [80]. In 2014, Serai et al. proved the technical feasibility of UTE-MRI in pediatric musculoskeleton at a clinical 1.5 T system with less than 5 min TA, supporting its beneficial use for the evaluation of bone pathology, such as irregular ossification at the weight-bearing site of the femoral condyle [81].

Another challenge in MR imaging of hard tissues is to image bone structures in technically difficult areas. For instance, MR imaging of the hip is considered technically challenging due to the presence of a thin cortex, the requirement of a robust suppression of the surrounding soft tissues, and the lack of commercially available dedicated coils. A 3D UTE-Cones sequence, which employs a hard RF pulse for non-selective signal excitation followed by a 3D cones trajectory for k-space sampling, allowed time-efficient sampling with a minimal nominal TE of 32 μs. The 3D UTE-Cones was combined either with an adiabatic inversion recovery (3D IR-UTE-Cones) or a dual inversion recovery (3D DIR-UTE-Cones) preparation pulse. Such techniques were used for the visualization of the cortical bone in the hip, tibia, knee, and ankle, and for the quantification of the associated T₂^* values (Table 2) [34, 82]. Furthermore, the combined 3D UTE-Cones sequence with a multi-spoke per MT preparation and with a modified rectangular pulse (RP) model allowed for a fast-volumetric quantification of the water content in short T₂ tissues by using a clinical 3 T whole-body scanner [83].

ZTE MRI was also applied for bone tissues on a clinical scanner and improved simultaneously along with the development of UTE methods.

Water- and fat-suppressed proton projection MRI (WASPI) represents a type of zero echo time pulse that allowed to excite only short T^*₂ components based on a T₂-selective RF excitation method consisting of long-duration low-power rectangular RF pulses [87]. MRI-WASPI was used to image the solid matrix content of rat femora with a spatial resolution of 0.4 mm, hence representing a useful tool for the quantification of the bone matrix density [88]. A further implementation of ZTE MRI for humans was reported in 2013 by Weiger et al. (Table 2) [35]. The authors reported the first 3D ZTE images of the human head, wrist, knee, and ankle on a 7 T human whole-body MRI system. The acquired images showed fine anatomical details with a well-delineated bone tissue interface, high SNR, and isotropic spatial resolution of 0.83 mm [35]. In addition, ZTE method was also used on a 7 T μMRI system to image trabecular microstructures in bovine bone samples, and the images were compared with μCT results. With respective imaging resolutions of 56 μm and 14.8 μm for ZTE and μCT acquisitions, the trabecular micro-architecture was shown with excellent agreement on both modalities. Furthermore, the assessed bone volume fraction resulted in similar values, i.e,. 0.34, and 0.36, for both ZTE imaging and μCT [89]. Afterwards, a rotating ultrafast imaging sequence (RUFIS) type of ZTE imaging (ZTE-RUFIS) was developed using a non-selective hard pulse excitation followed by a 3D center-out radial sampling (Table 2). With ZTE-RUFIS sequence image acquisition started immediately leading to a nominal TE equal to zero. Furthermore, by using minimal gradient switching in between repetitions and short RF pulses, a robust and fast method with more efficient SNR was achieved. Such ZTE-based method provided high-resolution pictures of the cranium, facial skeleton, and cervical vertebrae of human volunteers. MRI acquisitions depicted anatomical details equivalent to the CT acquisitions obtained in a PET/CT scanner [36].

Additionally, ZTE technique was associated with a MT method and tested for the imaging of the cortical bone composition in mice on a 4.7 T system (Table 2). Relaxation time properties of the femoral diaphysis that were based on MT-ZTE acquisitions were reported, and the values were 1107 ± 203 ms for T1, 12.5 ± 2.0 μs for T2, and 563 ± 75 μs for T2*, which resulted to be higher when compared to previous studies. These differences were not only ascribed to a difference in magnetic field strength that was different when compared to other studies, but also to the higher sensitivity of the MT-based approach [38].

A in vivo comparison study between UTE and ZTE methods was performed on a 7 T system [90]. After MRI acquisition of the bone in the brain, ankle, and knee of human volunteers the results dealing with the imaging quality, resolution capabilities, and off-resonance sensitivity were compared to each other. This study showed no differences in resolution capabilities and comparable image contrast and SNR for both UTE and ZTE images. Only subtle differences were found in the off-resonance response, due to the difference in k-space sampling for these two MRI techniques. ZTE acquisitions showed increased blurring around the skull, as well as signal dropout artifacts in proximity of the edge of the field of view, while UTE methods showed more flexibility in imaging volume selection resulting more adequate for clinical applications [90].

The use of MRI in dentistry is continuously increasing. Dental tissues, however, show the shortest relaxation profile, making their imaging the toughest challenge for MRI. The first UTE-MR image of teeth was acquired by Gatehouse et al. in 2003 [29] through FUTE method on a 1.5 T system. Afterwards, human volunteers were also scanned using a whole-body 3 T system with a total scanning time of 10 min [93]. The sagittal slice of the jaw acquired by UTE sequences with TE = 50 μs provided a clear view of all the dental structures including enamel (Table 3). Subsequently, 3D-UTE sequence was used not only to image teeth but also to estimate the relative T₂^* values, which could be correlated with the presence of dental caries [94, 96, 97]. In fact, the mineral breakdown caused by caries was found to lead to a gradual increase in T₂ value, depending on the extension of the lesion, hence supporting the use of UTE-based MRI as a feasible screening tool for the early detection of dental caries.

SWIFT sequences have shown the possibility to obtain simultaneous acquisition in vivo of both soft and hard dental tissues with high resolution and short acquisition time (Table 3) [94]. SWIFT-MRI acquisitions were performed on a 4 T system and with a custom-made one-side shielded coil, which was located between the cheek and the teeth. The shielded coil was able to minimize the signal from the surrounding soft tissue. Furthermore, to reduce artifacts associated with the patient motion, a single motion correction was performed by comparing 16 low-resolution images. Furthermore, SWIFT-MRI has been used for the ex vivo detection of cracks into human molars. The presence of higher water concentration into the cracks when compared to the surrounding dentin enables the detection of lesions up to 20 μm in width [98].

High-resolution ZTE images of extracted teeth (i.e., incisors, canine, molars, and wisdom teeth) were made ex vivo using an 11.7 T MRI system. MR images of the teeth showed various dental caries and dental fillers, including amalgam and ceramic materials. Three-dimensional ZTE images with 148 × 148 × 188 μm resolution were acquired in 6.55 min and compared with UTE-MRI and μCT acquisitions. ZTE-based MRI was able to highlight slight differences in dentin composition that were only barely or not recognizable at all by μCT. Furthermore, ZTE-MRI showed signal variation around the pulp, which was not visible in μCT, and local signal changes in the enamel and dentin. However, when compared to the ZTE-MRI, μCT showed better signal from enamel and dental calculus as present on the tooth surface. When UTE- and ZTE-based images were compared, ZTE imaging showed a better contrast of both dentin and enamel as also confirmed by the quantification of the SNR for each imaging modalities [95, 99].

Information about the mineral density of the teeth and of their major component (i.e., HA) was obtained by quantification of the phosphorous content. Extracted human molars were scanned on a 9.4 T system with a home-build double resonance (P-31 and H-1) probe through a ³¹P-SWIFT and ³¹P-ZTE sequence. These methods provided ³¹P-based images, which made it possible to distinguish the enamel from the dentin structures with a resolution of 0.5 mm. When comparing the two MRI methods, the P-31 MRI images obtained through ³¹P-ZTE imaging showed better SNR (> 28 %) than ³¹P-SWIFT [100].

The lack of coils and powerful gradient in a whole-body MRI system is hampering the translation of ZTE-MRI into the clinic [101]. An intraoral MRI coil that can be placed between the maxillary and mandibular teeth has recently been proposed by Idiyatullin et al. [102]. The coil consisted of a single loop of 10 mm width and covered with sticky foam to be suitable for the adult maxillary arch. This coil was used for MRI acquisition of a human volunteer at a 4 T MRI scanner through a 3D radial SWIFT sequence. The intraoral coil showed high SNR and spatial resolution (0.3 mm³) compared to an extraoral coil. However, the main limitation of the intraoral coil was the lower SNR in proximity of the cusps of the teeth [91]. Ludwig et al. developed a wireless inductively-coupled intraoral coil consisting of two coaxial loops of 1.5 and 2 cm diameter. The coil was covered with a dental resin and adapted to the human mouth anatomy by a home-made dental cast [103]. The coil was used firstly ex vivo on a porcine mandible and then in vivo in a human volunteer on a 3 T MRI system. However, instead of an ultrashort TE sequence, a high-resolution 3D-FLASH (fast low angle shot) was used for the image acquisitions. The coil showed improved SNR when compared to other dental coils described in the literature and displayed relevant anatomical details with an isotropic voxel size of 350 μm. The optimization and combination of the coil with ultrashort echo time sequences remain still unaccomplished [103].

The development of MRI sequences that can detect short T₂ components widens the use of MRI also for artificial materials, as used in orthopedic and dental applications, like bone fillers and restorative materials [104]. Materials that are used for bone and dental application can be distinguished based on their degradation properties in degradable and non-degradable materials. Degradable materials, such as ceramics and certain polymers, have the property to be removed from the body through in vivo degradation, while non-degradable materials, such as metals and resins, are biologically compatible but mostly need to be removed from the body with an additional operation (e.g., metallic fracture stabilization plate) or left in situ (e.g., permanent dental fillers) [105, 106].

S. Emid and J.H.N. Creyghton [107] firstly proposed in 1985 a single-point imaging (SPI) method, also known as contrast time imaging (CTI), to study solid materials with very short T₂ values (e.g., > 50 μs). SPI is a pure phase-encoding imaging method which uses a single data point in the k-space that is acquired after a short encoding time in presence of a gradient. Although SPI method has been widely used for special purposes, such as studying solids, removing susceptibility artifacts, and chemical shifts artifacts, long acquisition time was the main limitation. Conventional SPI has been recently implemented to enable continuous imaging with a more time-efficient acquisition (CSPI, continuous single-point imaging) [108].

In 2011 Springer et al. [109] proposed a modified Ernst equation and variable flip-angle method combined with a 3D-UTE sequence to visualize polymeric materials with a 3 T whole-body MRI scanner. The variable flip-angle method allowed for the T₁ quantification of the polymeric material within a reduced acquisition time. Although the exact materials composition was not described, the authors were able to estimate their relaxation properties, i.e., T₁ = 223.1 ms and T₂^* = 0.295 ms [109]. Dental restoration materials, such as amalgam and ceramic inlays, inserted in human molars have been imaged with high-resolution ZTE sequence and compared with μCT acquisitions. Dental fillings caused image artifacts in both modalities depending on their compositions. Specifically, amalgam filling, which was used in a molar after endodontic treatment, caused a strong beam-hardening artifact compromising the μCT imaging. Differently, in ZTE images of amalgam filling the artifact was reduced and confined to the filling location allowing for better diagnosis [95]. MR imaging artifacts are very common also in presence of metallic implants, such as hip prosthesis, due to the large magnetic susceptibility differences between the implant and the tissue, which results in perturbation of B₀. A fully phase-encoded UTE (UTE-FPE) method has been proved to give distortion-free images near the metallic implants in human subjects and in about 5 min acquisition time (Table 2) [33].

A great variety of dental restoration materials have been imaged by UTE-MRI thus far. For all these compositions the relaxation profiles (i.e., T₁ and T₂^* values) were quantified and are available for further optimization of MRI sequences (Table 4) [110].

Biodegradable calcium phosphate-based (CPC) compositions have also been imaged by MRI. Sun et al. [111] investigated the MRI visual properties of a specific injectable CPC composition consisting of a mix of 68 % alpha-tricalcium phosphate (α-TCP), 8 % dicalcium phosphate dehydrate, 4 % HA, and 20 % poly (lactic-co-glycolic acid) (PLGA). The CPC was first injected in vitro in a cylindrical defect created in bone blocks and subsequently implanted in vivo in cylindrical defects made in rat femora. The assessments were performed on an 11.7-T scanner with both UTE and ZTE sequences. Relaxation estimation for the CPC composition based on a 3D-UTE sequence showed a T₂^* equal to 442 μs. This value was very close to the T₂^* relaxation of the pig cortical bone (i.e., 597 μs), resulting in a similar MRI signal with both UTE and ZTE acquisition. This similarity in relaxation profiles hampered the identification of the CPC after implantation either in vitro or in vivo, which makes the use of MRI contrast agents a proper solution for the enhancement of the CPC signal [111]. Therefore, CPC has been combined with different contrast agents, such as superparamagnetic iron oxide particles (SPIO), gadolinium-based contrast agent (GBCA), and perfluorocarbons (PFC), aiming to a better visualization of the implant and a more precise in vivo follow-up [112‐114].

CPC combined with a dedicated dual contrast agent (DCA) consisting of SPIO with a mean size of 200 nm, and gold nanoparticles (AuNPs) with a mean size of 4 nm (respectively, as MRI and CT contrast agent) was imaged in vivo by ZTE-MRI and CT after installation in rat femoral defects. Longitudinal ZTE acquisitions allowed the identification of the implanted CPC until 8 weeks post-implantation, while the CT contrast of the CPC was lost 4 weeks after implantation [99]. However, because of the strong susceptibility properties of the SPIO particles, a strong imaging artifact (named blooming effect) was observed on the ZTE-MR images. The blooming effect leads to an easy identification and localization of the implanted CPC but hampered the morphological analysis of the material [112]. Interestingly, blooming artifacts are also observed after ZTE-MRI acquisition of CPC combined with GBCA nanoparticles, which are known to show a T₁-weigthed contrast agent [113]. The appearance of such artifacts could be partially ascribed to the high magnetic field strength (i.e., 11.7 T). Nevertheless, the use of fluorine-based contrast agents is considered as a possible solution to diminish imaging artifacts. Therefore, CPC was combined with polymeric nanoparticles containing PFC (i.e., perfluoro-15-crown-5-ether) and gold nanoparticles, respectively for ¹⁹F-MRI and CT imaging. Fluorine-based ZTE-MRI of the labeled CPC implanted in rat femora permitted the fine visualization of the CPC shape and allowed the monitoring of the material degradation longitudinally up to 8 weeks post-implantation [114].

Another CPC composition, consisting of 59.1 % α-TCP, 1.5 % carboxymethylcellulose (CMC), and 39.4 % cryo-grinded PLGA, was used as a dental pulp capping agent in extracted human molars. The specimens were scanned on an 11.7-T scanner through a 3D-UTE and ZTE sequence. This specific CPC reported a T₂^* equal to 273 μs, which was lower than the T₂^* from human dentin (i.e., T₂^* = 476 μs). Therefore, such a CPC can be distinguished from dentin on the ZTE acquisition (Fig. 7). Also, relaxation studies were performed on the CPC before and after 7 weeks implantation in goat incisors. A lower T₁ value was found for the CPC composition after the implantation in vivo, i.e., 742 ms versus 1008 ms after and before the in vivo implantation, respectively. These results suggested the feasibility to use the relaxation properties of the CPC to follow its degradation in vivo [115].

MRI sequences for the imaging of hard tissues (e.g., bone and teeth) and biomaterials used for restoration of such tissues are becoming more and more advanced. The huge potential of ultrashort TE MRI is evident, not only for the morphological high-resolution 3D reconstruction of bone and dental structures, but also for the early detection of bone and dental diseases. By using ultrashort TE MRI it is possible to obtain information about the status of a bone pathological condition. Moreover, MRI is considered a very promising screening tool in dentistry not only because of the lack of radiation, but also because of the simultaneous multiplanar imaging of soft and hard tissues as well as the ability to detect demineralization and dental caries [116].

However, many challenges still need to be solved before such MRI methods will be accessible to a wider patient population. While issues with hardware systems and fast transmit/receive switches can be solved in a relatively easy way, the need for high gradient performance remains still a challenge. One approach to image short T₂ component could be to use conventional MRI sequences and reducing the duration of the RF pulse to a value comparable with T₂ and using a readout gradient with much shorter time than the T₂. This approach is feasible in small-bore system but not suitable with whole-bore equipment. This is due to technical difficulty to build high-performing gradients (e.g., peak 40 T/m and slew rate of 2 T/m/ms) and RF systems capable to generate 10 μs 180° pulse (i.e., ˃ 500 kW RF amplified peak power) [14]. Furthermore, RF coils induce an electric field that results in deposition of energy into tissue in the form of heat. Such phenomenon can cause serious tissue damage especially in case of presence of metallic implants. Specific adsorption rate (SAR) limits have been set by regulations and are limiting the possibility to use higher-performing gradients [18, 117]. To date, safety regulations limit the maximum gradient switching to < 20 T/m/s, while RF excitation pulses are limited to < 4 W/kg [14]. However, it is important to mention that UTE sequences have less SAR limitations when compared to other ultrashort echo time techniques, such as ZTE and SWIFT, where RF is applied simultaneously with encoding gradients. Furthermore, the requirement of appropriate software, which allows for qualitative and quantitative analysis, also needs to be considered as not all MRI manufacturing companies updated these recently developed sequences [16]. Finally, the extensive costs for MRI machines and their maintenance still hamper a global spreading of such imaging tools, especially in dentistry. The development of a relatively “cheap” MRI system is necessary. This review reported many studies that can be used as a proof-of-concept for the translation of MRI to dentistry. Therefore, the next step will be the refinement of the available technologies to make them suitable for dental applications, while the use of short-bore systems and of low magnetic fields relying on the pre-polarization MRI concept can be suggested as a possible solution to reduce the manufacturing costs [118, 119].

In summary, ultrashort TE sequences represent an applicable MRI tool that allows quantitative and qualitative assessment of bone, teeth and solid-like biomaterials. Their clinical translation is an ongoing occurring process especially in the orthopedic field, while still many efforts need to be done in dental MR imaging. The use of biomaterials for bone and dental repair is well established in the clinic, which warrants the investigation of MRI sequences for their screening.