CROSS REFERENCE TO RELATED APPLICATION

[0001] This application claims priority to and the benefit of US Provisional Application No.

62/928,099 filed October 30, 2019, which is incorporated herein by reference in its entirety.

GOVERNMENT RIGHTS

[0002] This invention was made with government support under Grant Number R43CA232901, awarded by National Cancer Institute of the National Institute of Health. The government has certain rights in this invention.

TECHNICAL FIELD

[0003] This disclosure relates to medical imaging, more particularly to optical and MRI imaging using diamond particles.

BACKGROUND

[0004] In the quest towards high signal-to-noise (SN) imaging, the power that can be brought to bear by multimodal or multi-messenger techniques has been long recognized. At its heart such multimodal imaging entails capturing the object through more than one imaging mode simultaneously, often at widely disparate wavelengths. Exploiting correlations between the different channels portend Kalman filtering approaches that can deliver a high degree of noise or background suppression. Moreover, under appropriate conditions, these correlations engender novel image sampling and reconstruction strategies that can substantially accelerate image acquisition.

[0005] While routinely employed in astronomy, a number of these multimodal imaging techniques are now finding compelling applications in biomedicine. For instance, new instruments that combine magnetic resonance imaging (MRI) with radiotherapy have opened exciting new high-precision therapeutic avenues. Interest has gained ground for methods that similarly marry together MRI and optical imaging, with applications in precision 3D MRI- guided laser ablation, and in continuously MRI-monitored endoscopy for high-accessibility colorectal cancer screening.

[0006] Indeed, in several respects, MRI and optical imaging offer diametrically complementary advantages, a feature that can make them particularly powerful in combination. Visible-wavelength optics is fast, cheap and images at high-resolution; yet often suffers from scattering, attenuation, and aberration distortion when imaging through tissue. MRI, on the other hand, is noninvasive, fully three-dimensional and can be chemically functional; yet it is slow, suffers from weak signals, and offers poor spatial (mm-level) resolution. Perhaps more importantly, optical and MR imaging are carried out in Fourier- reciprocal spaces (x- and k-space). This redundancy immediately makes a combined modality highly persuasive. Not only are there complementary advantages in sensitivity, resolution, and aberration suppression to be gained, but also the possibility for hybrid new image acquisition strategies that sample both real and k-space simultaneously to yield meaningful imaging acceleration. If such dual-mode imaging is deployable in a biocompatible and functionally targetable material platform, several exciting avenues could result in disease monitoring as well as in targeted drug delivery.

BRIEF DESCRIPTION OF THE DRAWINGS [0007] Figure 1 A shows an embodiment of a dual -mode imaging apparatus.

[0008] Figure IB shows a graphical representation of the hyperpolarization protocol and detection.

[0009] Figure 1C shows a signal gain from a dual -mode imaging process. [0010] Figure ID shows an embodiment of a phantom ring filled with diamond particles used for dual-mode imaging.

[0011] Figure IE shows an embodiment of a phantom ring captured with optical imaging. [0012] Figure IF shows an embodiment of a phantom ring captured with magnetic resonance imaging.

[0013] Figure 2A shows a graphical representation of a normalized fluorescence signal for randomly oriented ensemble of diamond particles under an applied magnetic field.

[0014] Figure 2B shows signal contrast for optical modulation under a pulsed magnetic field. [0015] Figure 2C shows optical images captured when diamond particles were under a magnetic field.

[0016] Figure 2D shows examples of magnetic resonance images under opposite microwave field sweeps.

[0017] Figure 2E shows a graphical representation of the result of a positive magnetic field sweep and a negative microwave field sweeps.

[0018] Figures 3A-B show a graphical representation of an imaging die.

[0019] Figures 3C-D show examples of magnetic resonance and optical images with a strong artificial background.

[0020] Figures 3E-F show examples of magnetic resonance and optical images with the background suppressed.

[0021] Figures 4A-B shows an embodiment of a process of accelerated imaging with accompanying images.

[0022] Figure 4C shows a graph of sampling width versus normalized time versus acceleration.

[0023] Figure 4D shows a graph of scaling of optimal sampling sizes with scarcity. [0024] Figure 4E shows a graph of optimized imaging acceleration through sampling of an optimal number of points.

[0025] Figure 4F shows a graph of a trajectory of convergence between two images.

[0026] Figure 5 shows a schematics of the correlative MRI/optical colonoscopy using targeting fluorescent nanodiamond probes with ¹³ C optically hyperpolarized at room temperature.

[0027] Figure 6 shows a schematics of a method of dual mode imaging using hyperpolarized diamond particles.

DETAILED DESCRIPTION OF THE EMBODIMENTS [0028] The embodiments here demonstrate the gains made possible by dual mode imaging using diamond microparticles. In this discussion “dual mode imaging” means high contrast dual-mode optical and ¹³ C MRI imaging in diamond microparticles. These versatile materials have gained prominence for tumor targeting given their biocompatiblity and surface functionalizability with antibodies, dendrimers and other targeting ligands. In the embodiments, the diamond particles are incorporated with a large concentration (approximately > 1 ppm) of Nitrogen Vacancy (NV) defect centers.

[0029] Under sub-bandgap illumination <575 nm, the particles fluoresce brightly in the red with high luminosity (~90 cd/m ² ), with high optical stability, without blinking or bleaching. The NV centers are also endowed with attractive spin properties. The generation of fluorescence occurs concurrently with the optical polarization of the electron spins associated with them. This macroscopically large polarization (>10%) can be transferred to ¹³ C nuclei in the surrounding lattice, hyperpolarizing them in vastly athermal states that bequeath them with brightness in MR imaging. The embodiments here exploit a recent NV-mediated low- field (1-70 mT) room-temperature hyperpolarization technique that allows large ¹³ C polarization levels (~1%) and correspondingly high MR signal enhancement factors, over a factor of thousand at 1.5 T.

[0030] Indeed, these gains in MR solid imaging were traditionally possible only through the use of cryogenic conditions and high magnetic fields, and only at substantially lower throughput. Moreover, the ¹³ C polarization can be retained for long periods, approaching ~10 min for certain samples at moderate fields >100 mT.

[0031] Not only can the high SNR optical and MRI imaging proceed concurrently, in addition, there is the added advantage that signals from both modes can be modulated on- demand, allowing a high degree of background suppression through lock-in techniques. The embodiments demonstrate high contrast optical and MRI imaging in phantom samples, and show suppression of signal backgrounds by over factors of 2 and 5 respectively in the two imaging dimensions respectively. Crucially, the embodiments demonstrate how combined conjugate imaging in real and k-space can promote several orders of magnitudes in imaging acceleration, particularly in wide field-of-view scenarios. This work paves the way for high contrast, background-free, accelerated dual-mode imaging of biocompatible nanoparticle delivery and targeting agents based on quantum materials.

[0032] Fig. 1A shows a schematic of one embodiment of the experiment. Diamond particles such as 10 (200 pm, ~ 40 mg) arranged in a ring-shaped phantom, as shown in Fig. ID, are imaged optically under continuous 520 nm illumination from the laser 12 and 630 nm long- pass filtering, as shown in Fig. IE. The high intrinsic SNR from the optical florescence is evident, and the ultimate imaging resolution is just diffraction limited. For MR imaging, as shown in Figs. IB, and IF, the process employs dynamic nuclear polarization (DNP) at 38 mT 14 to enhance the ¹³ C polarization, and a microcoil 16 in a 9.4 T magnet for imaging. The same optical excitation polarizes (initializes to m _s =0) the NY electron spins, and microwave (MW) sweeps across the NV ESR spectrum drive Landau-Zener dynamics that transfers polarization to the ¹³ C nuclei in a fully orientation-independent manner. The process obtains ~0.3% spin polarization in 40 seconds under 1 W total optical illumination, and this is primarily laser power limited. At this level, this corresponds to a signal enhancement of 206 times over thermal ¹³ C polarization at 9.4 T and shown in Fig. 1C, ~5 X 10 ⁴ over 38 mT, and ~10 ⁵ acceleration in imaging time.

[0033] It is worth noting that hyperpolarization through these embodiments requires relatively low laser (-2 mW/mg) and MW power (~0.05 mW/mg), where the estimates are reported for equivalent mass-weighted SNR to images presented here. Considering the saturation regime for a single 1 pm diamond particle, it is estimated a 30 nW optical and 2 nW MW power requirement respectively. Due to the high mass of the particles here, the MRI experimental demonstrations were performed with laser power densities ~80 mW/mm ² , somewhat elevated above levels suitable for in-vivo operation (~4.7 mW/mm ² ); indeed, laser power is not restrictive in preclinical settings since substantially lower masses are typically employed. Moreover in the experiments here, it is estimated that a specific absorption rate (SAR) of 1.1 X 10 ⁴ W/kg[G] ² of MW power. At an electron Rabi frequency of «100 kHz, broadly defining the regime of the operation, this corresponds to «14 W/kg, which is close to the threshold of safe in-vivo operation (4 W/kg). One should note that the hyperpolarization signal is highly robust to decrease in MW power, and scales approximately logarithmically with it, indicating that the SAR can be easily curtailed while maintaining relatively large NMR (nuclear magnetic resonance) enhancement factors.

[0034] One embodiment uses a variant of FLASH (fast low angle shot MRI) to produce the MRI images at 9.4 T as can be seen in Fig. 1F. To overcome the short T2 ≈1 ms of ¹³ C in diamond, the width of the radio frequency (RF) pulses and the gradient lobes are minimized to create a short echo time (0.5 ms). Moreover, to eliminate any interference with the phase encoding gradient and distortions during the short RF pulses due to the relatively long (200 μs) gradient switching periods, the imaging was performed without a slice selection gradient as shown in Fig. IB. This is also possible since the primary interest lies in projection images along the z-axis.

[0035] The SNR of the MR image shown in Fig. IF is ~4 in 16 scans, limited by rapid ¹³ C T2 decay, low sample filling factor (≈0.0016) and laser-limited hyperpolarization. The use of line-narrowing sequences, such as spin-locking or quadratic echoes, can improve the imaging SNR by at least an order of magnitude. For in-vitro samples, the use of higher laser power close to saturation intensity (~1 W/mm3) can increase the MR signal 10 times compared with the present results. Diamonds with higher ¹³ C content will provide larger signals, with the resulting SNR improvement scaling with enrichment. Further improvements may be realized by optimizing the detection coil geometry and increasing of filling factor, for instance through the use of small volume inductively-coupled receiver coils matched to the sample under study. These concerted gains in MR signal (~3 orders of magnitude) could also permit similar high contrast images in nanosized particles (<100 nm), material platforms that are most germane to in-vivo applications.

[0036] The MRI spatial resolution scales a 1 /yG _max τ), where g is the gyromagnetic ratio, G _max and τ are the maximum gradient strength and the duration of its application. In these experiments, Gmax = 950 mT/m, leading to a resolution of 640 pm in both dimensions. After zero-filling and smoothening, the pixel size presented in the image shown in Fig. IF has a square length of 160 pm. Recent amplifier development that increase Gmax can improve the spatial resolution. Of course, ultimately, the MR resolution can be just optical diffraction- limited, since only illuminated diamond particles contribute to any signal. One can envision beam-rastering modalities that buildup MR images pixel-by-pixel, in a manner that is solely resolution limited by optics.

[0037] The optical DNP method discussed here presents several advantages when compared with traditional methods of hyperpolarization for solids imaging, employed for instance in ²⁹ Si microparticles. The process works at room temperature and low field (~40 mT), and polarize samples in under 1 minute of laser pumping, as opposed to traditional high magnetic field (approximately > 3T) and low temperature (< 4K) approaches where polarization buildup can take several hours. While the absolute polarization is lower in the methods here, they circumvent the traditionally high polarization loss, which can be as large as 99%, accrued upon thawing and sample transfer out of the cryostat; ultimately, this results in a relatively high level of polarization delivery at the imaging source.

[0038] Technologically too, the techniques here aid end-user operation. The MW amplifiers and sweep sources are low-cost and readily available, empowering the construction of highly portable hyperpolarization devices that can retrofit to existing MRI scanners. More importantly, since the high polarization is detection-field agnostic, the embodiments here would open possibilities for continuous low-field MRI, along with simultaneous optical imaging in-situ, particularly attractive because the hyperpolarization can be replenished optically.

[0039] Both optical and MRI modalities allow on-demand signal amplitude modulation, enabling common-mode suppression of background signals in both imaging dimensions. The fluorescence emission at moderately low fields, especially approaching 50 mT, is strongly conditioned on the misalignment angle θ of the N-to-V axis to the applied field. This arises due of the mixing of the m _s = ±1 spin levels in the excited state, and since the randomly oriented particles sample all possible # angles, allows a simple means to modulate the optical images by the pulsed application of an external field B _ext . The embodiments simulated the fluorescence dependence under application of B _ext , shown in Fig. 2A (i) on the left side, using rate equations in a 7-level model of the NV center: where n _i is the population of the \i> state, and k,, denotes the kinetic transition rate between state [0040] At steady state, the state population and consequent photoluminescence, shown as the red line in Fig. 2A, obtained evaluating decreases with B _ext , in reasonable agreement with the (normalized) experimental measurements, shown as the blue dots in Fig. 2B. In the imaging experiments here, the process modulates the optical images shown in Figs. 2C and the MRI images in Fig. 2D, by application of B _ext ~40 ±2 mT, an identical field employed for hyperpolarization, and obtain a 10% optical modulation contrast show in Fig. 2B.

[0041] MRI amplitude modulation relies on the remarkable dependence of the ¹³ C hyperpolarization sign on the direction of MW sweeps, a feature originating in the rotating frame LZ dynamics excited by the chirped microwaves as shown in Fig. 2E. The ¹³ C nuclei are aligned with the polarization field under low-to-high frequency sweeps, or anti-aligned with the high to low frequency sweeps, over the NV ESR spectrum. Importantly, and contrary to optical modulation, this allows a complete sign-reversal of the MRI images at full contrast, all while requiring no additional infrastructure. Indeed, it is challenging to achieve such high signal modulation contrasts with conventional cryogenic DNP approaches, primarily due to technical limitations of MW cavity switching in these experiments. As a figure of merit, the embodiments here characterize the modulation contrast as the difference ratio of the MR images under opposite MW sweeps positive and negative images as shown in Figs. 2E and 2F, as: where N ² is the total number of pixels. From the data in Figs. 2E one can obtain D =194 ± 0.3%. The difference from full contrast seems to arise from the imprecise repeatability of the pneumatic sample shuttler, an easily improvable feature, and irrelevant if the MR images are ultimately to be acquired at low-field. It has been demonstrated that D >198% for the hyperpolarized NMR signals.

[0042] Such high-contrast signal modulation opens the door to imaging the diamond particles with high SNR even while embedded in an overwhelmingly strong background. The term background in this context refers to media that have fluorescence or ¹³ C NMR signals that overlap in wavelength, or NMR frequency, with the diamond particles. As shown in Fig. 3, the process used particles being co-included with a high concentration of Alexa 647, a fluorescence dye which has a strong emission at 600 nm, as well as [ ¹³ C] -methanol, which has a chemical shift nearly overlapping that of diamond. These solution media fill both the inner and outer spaces of the capillary tube that comprise the diamond phantom in Figs. 3A- B. The strong backgrounds result in images that are circle-shaped since the diamond phantom is completely indiscernible within it as shown in Fig. 3C-D.

[0043] To recover the diamond signals in the optical modality, the process performed lock-in detection under application of a 40 mT 0.1Hz square-wave modulated magnetic field. A 2000-frame movie was recorded at 0.1 s frame rate, and computationally applied lock-in suppression on each pixel individually. The resulting image shown in Fig. 3E shows good recovery of the diamond signal, in this case, from the 2 times larger background. Ultimately optical background suppression is limited by the optical contrast, and up to an order of magnitude is routinely experimentally feasible. Concurrent MRI background suppression is realized by subtracting the images under opposite sweep-ramp hyperpolarization conditions. [0044] Although the [ ¹³ C] -methanol signal herein was 5 times larger than the diamond signal, it is completely canceled as shown in Fig. 3F. Indeed, hyperpolarization sign control allows one to address the diamond ¹³ C nuclei in exclusion of all other ¹³ C spins in the sample. The intrinsically high contrast of the diamond MRI sign-reversals, along with the low inherent tissue backgrounds at low-fields, augurs well for high-contrast in-situ MRI detection.

[0045] The ability to co-register images in both optical and MR domains can yield SN enhancements, especially while imaging in-tissue settings. Optical fluorescence imaging requires the ability to illuminate the diamonds as well as collect the resulting fluorescence. While the diamond particles are quite optically bright, imaging through tissue presents a set of challenges that often dramatically restrict the proportion of photons that can be usably collected. One must contend with round-trip attenuation and scattering losses (scaling ~ = λ ^-4 ). Additionally, the unusually high refractive index differential between diamond and its environment, and geometric solid angle constraints of the detector at high fields-of-view (FOVs) severely restrict the ultimate flux of collectable fluorescent photons.

[0046] At 650 nm through human tissue, for instance, attenuation and scattering lead to exponential signal losses with coefficients 0.1 mm ^-1 and 1.1 mm ^-1 , respectively. An order of magnitude calculation of the chain of losses from illumination-to detection is quite revealing. Green-to-red photon conversion is rather inefficient, ≥ 10 ^-5 even at relatively low NV concentrations 0.1 ppm. Total internal reflection limits photon transmission and retrieval through the randomly oriented particles by a factor ~10 ^-3 . Round-trip attenuation losses cause signal suppression by as much as -0.3 through 5 mm of tissue. Scattering affects the signal by an excess of - 10 ^'3 for the same tissue thickness. Finally, finite numerical aperture detection limits photon collection by a further -10 ^-2 , here assuming a detection numerical aperture NA = 0.18. Put together, these losses can be considerable, and proscribe the ultimate efficiency of optical imaging in real tissue media. [0047] An analogous analysis of hyperpolarized MR imaging indicates that it can be competitively efficient, especially with increasing tissue depth. In contrast to optical imaging, one only partakes of losses in the one-way illumination of the particles with 520 nm light. Hyperpolarization efficiency is ~2 x 10 ^-4 per ¹³ C nucleus per incident photon. The high detection losses and geometric solid-angle collection constraints are replaced by more benign factors related to sample-coil filling, detector Q, and the overall MR detection frequency. Surface coils matched to the sample and the use of high-Q ferrite resonators can lead to substantially efficient detection. Indeed given the immunity to losses upon fluorescence scattering and attenuation, MR imaging can become preeminent even for buried particles relatively shallow in depth. If h were to be the ratio of optical and MR imaging SNR for surface diamond particles, MR imaging would have higher overall SNR than its optical counterpart at a depth d ~ — log(η) mm at 650nm, scaling logarithmically with the differential SNR ratio. Importantly, since both optical and MR imaging can proceed simultaneously, there are opportunities to use a “maximum-likelihood” hybrid of both modalities, wherein MR gains prominence with increasing depth profile.

[0048] Apart from reaping higher SNR, the ability of optical and MR imaging to capture the object simultaneously in Fourier-reciprocal spaces opens possibilities for significant imaging acceleration. As a specific example, one should consider atypical wide field-of-view (FOV) imaging scenario where optical imaging is performed by rastering a low power beam across the sample. This is germane to the imaging of embedded diamond particles in tissue. Simultaneous dual-mode imaging in real-space and k-space naturally provides the advantage that every sampled point in one space contains information from all the points in the other space. In the limit of imaging sparse objects in a wide FOV, and assuming that the imaging time cost is close to identical in both spaces, one could devise a hybrid protocol for sampling the FOV in both spaces simultaneously such that one rapidly converges to a high-fidelity image. Philosophically, this approach shares similarities with acceleration attainable through compressed sensing methods. In contrast, here the embodiments harness the ability to directly sample in both Fourier-conjugate spaces at once.

[0049] Fig. 4A describes one such imaging protocol. Given a sparse original image to be acquired, a subset l k-space points are first sampled in each dimension. The resulting blurry low-k image is thresholded and fed-forward to confine real-space points to be sampled through optical means. At high sparsity s, defined here as the fraction of zero pixels in the FOV, this can substantially reduce the number of points in real-space to be rastered over, and accelerate image acquisition. Fig. 4B presents an illustrative example of the protocol applied to an image with high sparsity, (1— s) = 0.0375. By employing £=16 k-space samples in each dimension, and thresholding at 75% of the maximum, one obtains an imaging acceleration of 16, since large swathes of the (real-space) image do not need to be sampled over. It is envisioned that this hybrid mode will be relevant to tumor-targeted diamond particles in-vivo, especially in endoscopic detection settings, where low-k MRI can serve as a guide to optical imaging.

[0050] The discussion can now analyze the regime of applicability and inherent trade-offs involved. In experimental settings, there is a time-cost to be accrued per sample (pixel) for both real-space (optical) as well as k-space (MR) imaging. For simplicity, it is assumed here that this time cost is identical for both imaging dimensions, although it is straightforward to scale the obtained results by the cost-ratio h as appropriate to the particular imaging setting. Given this, it is not surprising that there is an optimal k-space sample l _opt that draws the compromise between better constraining real-space sampling, and taking longer to determine. This is demonstrated in Fig. 4C where, in a 32 x 32 pixel FOV, one can see the normalized imaging time savings over either modality for target images with varying sparsity. In these simulations, the process performed a statistical analysis with 30 image configurations of fixed sparsity, assuming that the minimum feature size occupies one pixel. Fig. 4C demonstrates that hybrid sampling can deliver more than an order of magnitude in time savings (right axis), while only requiring the scanning of l _0pt ≈10% of total k-space (upper axis), notwithstanding the relatively small FOV considered. As expected, l _opt increases with decreasing sparsity, a reflection of larger k-samples required to account for the increasing image complexity, see Fig. 4D. Given the small FOV and discrete values of k-samples, this manifests in the staircase-like pattern in Fig. 4D, but scales l _opt α (1— s) ^1/4 (solid line) as derived below. Finally, Fig. 4E shows the combined imaging time t under optimized conditions as a function of image sparsity, assuming that time for optical imaging is 1. Indeed, the imaging acceleration is quite substantial, scaling as T ¹ α (1 s) ^-1/2 , and becoming increasingly utilitarian at high image sparsity.

[0051] It is revealing to consider the origins of this imaging acceleration by studying the trajectory of the reconstructed image as it approaches the target with each step of the protocol advancing, as can be seen in Fig. 4F. Once again studying several image configurations with a fixed sparsity, Fig. 4F shows the overlap of the reconstructed image I' to the target image / through the correlation C = Σ(l — (/)) (/' — (/')) where (·) indicates the mean value. Indeed under usual rastered form of optical image acquisition, shown as a dashed green line in Fig. 4F, the reconstructed image linearly approaches the target as more samples are acquired. In contrast, employing the hybrid acquisition of a few k-space points, and by constraining the space over which the final image is to be acquired, one obtains a rapid convergence of the image with the target. Numerically, the slope of convergence scales approximately α (1 — s) ^-1/2 shown as the inset in Fig. 4F, indicating rapid gains can be amassed at high image sparsity.

[0052] To analytically elucidate the imaging acceleration gains at high sparsity, show as the solid lines in Figs. 4D-E, one looks at the real-space target image f(x, y), which in k-space is where F denotes a spatial Fourier transform. As k-space sampling now occurs just to -f-th order, one obtains the reconstructed image, where II is a rectangular function representing sampling window with a side length of, for instance, W _kx = δk _x l where δk _x = 1/N _x is the k-space pixel size. Transformation back to real-space gives the convolution, F Indeed considering the imaging point spread function f(x,y) = δ(x,y), and assuming for simplicity a square image FOV, one obtains Hence an

N object of pixel radius r is effectively blurred where the factor ro is set by the thresholding level employed, and l/N is the effective k-space sampling ratio. Indeed, the k- space imaging constrains the area required for subsequent optical scanning to just this blurred region; increasing l makes a more faithful representation and improves regional constraints, but comes is associated with a time-cost. To evaluate the time savings as a result of this hybrid strategy, let us for simplicity now consider the FOV consists of n _d objects of radius r, giving (1 — s) = n _d π _r 2 /N ² . The normalized imaging time is then, In the limit of high sparsity, (1 — s)/r ® 0, and (1 — s)/r ² → 0(1), giving is a constant. Determining the optimal l to minimize t, gives l _opt α (1 — s) ^1/4 , and the optimal (normalized) time t α (1 — s) ^1/2 . The scaling from this simple model is shown as the solid lines in Fig. 4D-E and are a close match to the numerical results especially at the limit of high image sparsity. Here the image is considered to consist of unit pixel objects (2 r = 1), with no constraints on their relative placement in the image FOV. Moreover, while the discussion have restricted itself to 2D images, identical acceleration gains area attainable even for 3D imaging.

[0053] In summary therefore, leveraging Fourier conjugate sampling, even a simple protocol like the one in Fig. 4A can deliver remarkable gains at high image sparsity (s ® 1). The combined image acceleration scales t ^-1 α (1 — s) ^-1/2 , while only requiring the scanning of a minuscule percentage of k-space l _opt α (1 — s) ^1/4 . and leading to an approach to the target image whose slope scales oc (1 — s) ^-1/2 with each step of the protocol. Indeed, since this scenario is strongly operational in real-world targeted nanodiamond imaging applications, where sparsity factors s >95% are typical, it is anticipated that these strong acceleration gains are well within reach in practical settings. Since in principle the optical imaging can be performed during MR imaging, several possibilities exist for the protocol in Fig. 4 itself be further augmented.

[0054] The embodiments here have demonstrated a new means of dual-mode imaging in diamond microparticles, and shown that a slew of complimentary advantages can be harnessed by marrying together visible optical and ¹³ C MR imaging. The biocompatible particles employed have a high density of NV defect centers. The same optical excitation that causes high-luminosity visible fluorescence from the NVs also serves to spin-polarize the lattice ¹³ C nuclei making the particles light up in MR imaging. The embodiments demonstrated MR signal enhancements by -204 over 9.4 T under weak (-80 mW/mm ² ) optical illumination, and simultaneously bright optical florescence. They also demonstrated the ability to modulate the signals on-demand from both modalities, enabling a high-degree of background suppression in both optical and MR dimensions. Co-registering optical and MR images portends high-SNR imaging, especially for samples embedded in tissue and other scattering media, whereupon the MR modality can prove competitively efficient since it is largely immune to photon detection losses. Indeed, some embodiments involve MR imaging only, as the MR images do not suffer from the photon loss, and the improved efficiency and imaging due to the hyperpoloarized diamond particles.

[0055] Importantly, imaging in optical and MRI domains inherently occurs in Fourier- conjugate spaces, and allows development of hybrid protocols that feed-forward information from one domain to another to vastly accelerate image acquisition. For pertinent scenarios of imaging targeted particles in-vivo, where the wide-field images are highly sparse, the potential to accelerate image acquisition by about two orders of magnitude has been demonstrated.

[0056] Several straightforward pathways exist through which imaging SNR in both modes can be enhanced further. In particular, materials advances would boost MR image SNR by two orders of magnitude through ¹³ C enrichment and other sample processing that improve the hyperpolarizability of the diamond particles. Especially exciting amongst them are new methods for diamond rapid thermal annealing (RTA) at elevated temperatures (-1700-1750 °C) that have been shown to substantially boost hyperpolarization enhancement levels (in some cases by an order of magnitude), as well as ¹³ C T ₁ times, by suppressing deleterious lattice paramagnetic impurities. Such MR and optical imaging in-situ at low field might open exciting avenues for continuously optically replenishable imaging across both dimensions and might serve as a valuable targeting tool for several diagnostic and therapeutic avenues. [0057] Finally the high surface area diamond particles present an attractive platform through which the hyperpolarization could be transferred out into the targeting groups, including but not limited to antibodies, cellular receptors, proteins, DNA, drug molecules, metabolites, pyruvate, free radicals, intracellular components, cellular membranes, extracellular components and other biological substances potentially making them chemically-functional dual-mode imaging agents. Hyperpolarization transfer from diamond particles can be also considered for the targeting groups comprising non-biological substance.

[0058] With this system is it possible to consider a particularly compelling application for dual-mode imaging diamond particles in MRI-guided endoscopy during colonoscopy. Estimates show that colonoscopy misses up to 22% of adenomas, irrespective of size, and up to 2% of adenomas equal to or greater than 10 mm. While targeted fluorescence molecular imaging (FMI) has been suggested as a means to detect premalignant neoplasms and cancerous lesions in colorectal cancer screening, it cannot provide sufficient resolution for a signal collected from mm-deep location nor a complete 3D picture of the tumor. Instead, targeting nanodiamonds can label malignant lesions spread deeper into tissue. A simple optical fiber (power ≈10mW) fitted on the endoscope would serve to illuminate these particles in-vivo, enabling their imaging directly optically through the endoscope, as well as through low-field MRI.

[0059] Given the high inherent 3D FOV possible through MRI, malignancy-maps obtained from MRI could potentially aid in on-site optical endoscopy examination and lower experiential dependence, which currently plays an important role in the lesion detection rate. Moreover, by employing hyperpolarized diamond particles as major elements in MRI-guided needle tracking during, for e.g. biopsy of deep organs, would increase a visibility of needles and/or catheters in MRI images. Visualization of “passive” needles is based on their materials properties and is prone to imaging artefacts and inadequate tracking, while a recently emerging “active” MRI catheter devices are based on incorporation of fiducial markers along the needle profile to accomplish greater MRI visibility. Modulation of fluorescent and MRI signals from diamond particles open perspectives for completely background free tracking of catheters devices in-vivo.

[0060] Experiments were performed with -200 pm diamond powders (average edge length 87+3.9pm) with -1 ppm NV from Element6. While these particles provided high hyperpolarization enhancements (-280 at 7 T), hyperpolarization enhancements factors of -3 for lOOnm particles containing -3 ppm of NV centers were observed. A detailed discussion of material factors affecting ¹³ C hyperpolarization is beyond the scope of this discussion. The mass of the sample used in all the experiments is -40 pg. For uniform illumination of the high mass samples the optical imaging is performed under simultaneous excitation from four 520 nm fiber-coupled laser diodes (Lasertack) in a rhomboidal configuration, and the florescence imaged on a CMOS detector (Thorlabs DCC1645C) through a 594 nm long pass filter (Semrock BLP01-594R-25).

[0061] The hyperpolarization apparatus contains laser excitation, microwave irradiation and weak electro-magnet to fine-tune the polarization field. The embodiments employ a miniature 1W 520 nm diode laser (Lasertack PD-01289) in a feedback loop with an integrated thermoelectric cooler for adequate thermal control (TE Inc. TE- 63-1.0-1.3). Microwaves are generated by miniature voltage controlled oscillator (VCO) sources (Minicircuits ZX95-3800 A+, 1.9-3.7 GHz, output power p = 3.1 dBm). Frequency sweeps are produced by controlling the VCO frequency by a homebuilt quad channel voltage ramp generator controlled by a PIC microprocessor (PIC30F2020). The sweep generator employs dual multiplying digital-to- analog convertors (MDACs, Linear Technology LTC1590) to generate the sawtooth voltage ramps. All MR images are taken with the particles immersed in water or solution.

[0062] A pneumatic field-cycling device was implement to enable rapid sample transfer from low field (40 mT) to high a wide-bore 9.4 T Bruker DRX MRI machine, within which a 10 mm 1H/ ¹³ C Volume Coil is installed. The shuttling device is composed of a quartz channel transporting the sample, a concave-shape stopper at the end of the channel, and driven by a vacuum machine to transfer the sample in under Is.

[0063] Dual-mode imaging technology using hyperpolarized fluorescent diamond particles is particularly useful in clinical settings involving endoscopic imaging, including but not limited to the screening, surveillance, maintenance, or management of colorectal cancer (CRC), bladder cancer, gastro-Intestinal (GI) cancer, and other pathologies where endoscopy is applicable. Other clinical settings where the dual-mode imaging can be of particular use include but are not limited to prostate cancer, sentinel lymph node detection, image-guided surgery, and detection and maintenance of skin cancer. In these setting depths of light penetration through tissue allows for hyperpolarization of diamond particles incorporated into the tissue and body fluids. Yet other targets for visualization using hyperpolarized diamond particles comprise blood clots, thrombi, including their degradation products, as well as other cardiovascular malignancies. The method of imaging using hyperpolarized diamond particles also include targets for visualization such as at least one of a stent, needle, biopsy needle, catheter, implant, where diamond particles can be incorporated into coatings of the medical devices and provide their visualization in a body using MRI or MRI/optical dual-mode imaging.

[0064] Particularly beneficial is MRI/optical dual mode imaging for colon examination. Colonoscopy is the gold standard to reduce risk by detection of both premalignant neoplasms and carcinomas. However, 86% of all post-colonoscopy CRC can be explained by missed lesions which could have been prevented. Though targeted-fluorescence imaging could improve detection accuracy by generating cancer-specific contrast not possible in standard white-light endoscopy, the colon’s geometric complexity is an encumbrance which limits optical detection. The use of MRI to inform the endoscopist pre- or intraoperatively could greatly reduce miss rates; however, sensitivity is not enough to detect small lesions. The dualmode imaging technology of the present invention based on safe targeting diamond particles contrast agents endowed with bright fluorescence and strong nuclear magnetic resonance ¹³ C signatures can significantly increase fidelity of the detection and improve the detection rate of precancerous and cancerous lesions in colorectal cancer. Leveraging the MRI and optical detection of labeled malignances during colonoscopy, performed either sequentially or concurrently, could greatly improve the sensitivity of lesion detection. MRI provides a map of the lesion distribution within the colon informing the correlated optical examination, while fluorescence-guided colonoscopy highlights lesions on the colon surface to facilitate adenoma removal and biopsy as currently practiced and example of which is shown in Fig. 5. [0065] In Fig. 5, the hyperpolarized diamond particles 20 have been attached to a targeting ligand 22. These are then inserted or otherwise placed into the body where a malignant lesion such as 24 may exist. The contrast regions such as 26 result from the diamond particles seen by the MRI, in the MRI image of the colon 28. The luminescence within the colon is seen at 30, with the hyperpolarizing device 32 allowing optical imaging of the diamond particles, seen in fluorescence imaging 34.

[0066] Employed concurrently, both modes harness complementary advantages in sensitivity, resolution, and aberration suppression relevant to biomedical use. Dual-mode MRI/fluorescence endoscopy allows imaging across multiple length scales, with potentially cellular level resolution optically and organ-level information via MRI as shown in Fig. 5.

This cross examination by two modalities can lower error due to human factors, which currently plays a role in the detection rate. Importantly, based on strong optically hyperpolarized ¹³ C NMR signals, the technology can bring MRI’s detection limit of labeled lesions to 1-2 mm and potentially to 0.5 mm, advancing past the current capability limited to 6 mm lesions.

[0067] The dual mode imaging technology couples fluorescence colonoscopy and low field MRI respectively comprising two essential components: targeting of cancer ous/precancerous tissue in the colon with the fluorescent contrast reagent and the technique of the in vivo reagent hyperpolarization, where a simple optical fiber and a low power microwave source fitted on the endoscope would serve to illuminate these particles in vivo, enabling their imaging directly optically through the endoscope, as well as through low-field MRI.

[0068] One embodiment involves method of dual mode imaging using hyperpolarized diamond particles as shown in Fig. 6. At 100 the process provides diamond particles with enhanced hyperpolarizability. These diamond particles then have targeting ligands attached to them at 102. The ligands may be selected based upon their ability to increase cellular uptake of the diamond articles.

[0069] The process then attaches the diamond particles with the attached ligands to a target at 104 and removes unbound particles at 106. The diamond particles then undergo hyperpolarization at 108 by application of a specific sequence of microwaves and magnetic field while being illuminated with light. The process then images the target with attached hyperpolarized diamond particles using a magnetic resonance imager at 110, while also imaging the target using fluorescence imager at 112, although these processes do not have to occur simultaneously. The images from the MIR and the fluorescent imager undergo correlation at 114, resulting in a combined image. The process of hyperpolarization at 108 through the correlation at 114 may be repeated as necessary as shown at 116.

[0070] In addition to the repetitions and the simultaneously imaging, other variations and modification may exist. In one embodiment removal of the unbound particles can be omitted and a control diamond particles can be administered. Non-specific binding of particles to tissue may occur, and even washing may not remove them. Control particles help alleviate this issue and eliminate the need for washing. Control particles are particles emitting another fluorescent color but have not targeting ligands attached, but are administered at similar concentrations as targeting particles, those having ligands attached. If no tumor is present, particles of both colors will be randomly distributed in a similar manner. If problematic tissue is present, a significant excess of the targeting particles will attach to the cancer lesions over the non-targeting particles of the other color.

[0071] In one embodiment the diamond particles hyperpolarization may occur using an endoscope while imaged in a MR imager only, as discussed above. According to other embodiments, the target is at least one of cancerous lesion, cancerous cell, cancerous tissue, pre-cancerous lesion, neoplastic lesion, cell receptors, cancer markers, colorectal cancer, bladder cancer, GI cancer, prostate cancer, sentinel lymph node, cancer containing target in image guided surgery; skin cancer. In yet another embodiment the presence of a target confirmed in both MR imaging and fluorescence imaging increases the fidelity of a conclusion of the target’s presence and size.

[0072] One embodiment may include a method of imaging using hyperpolarized diamond particles comprises particles with size varied between about 10 nm and about 10 um; and where upon hyperpolarization the NMR signal enhancement varies between 0.1 and 20000 at 7 T. The diamond particles with enhanced hyperpolarization can contain NV centers in the concentration between approximately 1 ppm and 10 ppm. The hyperpolarized diamond particles provide fluorescence emission in at least one of NIR, red, green, blue, and UV spectral ranges and where illumination of the diamond particles with light comprises at least one of NIR, red, green, blue and UV spectral ranges, and the blue and green spectral ranges. [0073] In another embodiment, the attachment of the targeting ligands is done by chemical conjugation of at least one of an antibody, protein, peptide, molecule, folic acid, carbohydrate, or nucleic acid. An example of a targeted biomarker for CRC screening comprises epithelial growth factor receptor (EGFR), which is overexpressed in carcinomas. Targeting can be performed by conjugation fluorescent diamond particles (FDP) with approved antibodies, such as cetuximab and panitumumab. One embodiment includes diamond particles functionalized with poly(glycerol) by ring opening polymerization in neat glycidol to introduce colloidal stability in biological media followed by conjugation to cetuximab by carbodiimide or related activation. Alternatively, cetuximab fragments or other anti-EGFR antibodies with a stronger binding constant can be used. Additionally, biological compatible excipients can be used to stabilize the colloid including serum albumin and trehalose. Other targets of CRC comprise carcinoembryonic antigen, vascular endothelial growth factor A (VEGFA), or Integrin anb3.

[0074] One embodiment includes a method of imaging using hyperpolarized diamond particles comprises administering the diamond particles with the targeting ligands to a target comprising at least one of instilling with fluid into a body cavity; intravenous injection; locally; orally; systemically; intramuscular injection; intravascular injection; and intranasally. The removing of unbound diamond particles including but are not limiting to CRC imaging comprises instilling body cavities with liquid, removing liquid and repeating the process at least once. In another embodiment the unbound particles can be removed by a flow of body fluids.

[0075] Another embodiment includes the method of imaging using hyperpolarized diamond particles, where the hyperpolarization procedure comprises excitation with light from the spectral range of about 450 nm to 600 nm, preferably with excitation from 500 nm to 570 nm, and power between 1 mW and 100 mW; use of microwaves at frequency 0 - 100 GHz and power 1 mW-100 W; and magnetic field 1 mT to 3 T. The source of light, microwaves and magnetic field can be a part of a hyperpolarization device with dimensions and geometry suitable for insertion into at least one of a body cavity, tissue, organ and body fluid system. In one embodiment the hyperpolarization device comprises at least one of an endoscope, a colonoscope, acystoscope, and a laparoscope. In another embodiment it can be a free standing device.

[0076] Fluorescence imaging of a target with attached hyperpolarized diamond particles using fluorescence imager comprises using optical fibers inserted into at least one of a body cavity, tissue, organ and a body fluid, where imaging of the target with attached hyperpolarized diamond particles further comprises collection of fluorescence images taken in the presence and absence of a magnetic field and further comprises the images subtraction and producing a background free image of the target.

[0077] Another embodiment includes the method of imaging using hyperpolarized diamond particles, where imaging the target with attached hyperpolarized diamond particles using a magnetic resonance imager comprises low field MR imager; high field MR imager; NMR spectroscope; image modulation; image subtraction and lock-in procedure.

[0078] It will be appreciated that variants of the above-disclosed and other features and functions, or alternatives thereof, may be combined into many other different systems or applications. Various presently unforeseen or unanticipated alternatives, modifications, variations, or improvements therein may be subsequently made by those skilled in the art which are also intended to be encompassed by the embodiments.