Currently under study

The unique optical properties of gold nanoparticles make them attractive for a wide range of applications which require optical detection and manipulation techniques. Here, we experimentally demonstrate the use of single femtosecond pulses at resonance wavelength for a controlled conjugation of gold nanoparticles and fluorescent proteins. This optically driven reaction is rigorously studied and analyzed using a variety of experimental techniques, and a detailed model is proposed which describes the adsorption of the proteins onto the nanoparticles’ surface, as well as their subsequent desorption by a reducing agent. Potential applications of the resulting nanoparticle-protein conjugates include controlled delivery of fluorescent markers and local sensing of biochemical processes.

Figure 1

Fig. 1. (a) Fluorescence images of the nanoparticle-protein solutions following resonant illumination by a single pulse. The (false) colors correspond to the fluorescence emission wavelength of each protein. The field of view of each frame is 2.5 mm × 1.87 mm. (b) A bar chart showing the relative changes of the total fluorescence intensities after laser irradiation for protein-only solutions (blue bars) and for nanoparticle-protein solutions (green bars).

Figure 2

Fig. 2. Extinction spectra of illuminated nanoparticles solution with GFP before (dashed blue curve) and after a single pulse illumination (solid blue curve). Loss of the absorption peak after single pulse irradiation was observed when the nanoparticles were mixed with Myoglobin (solid black curve), Streptavidin (solid red curve), and without proteins (solid green curve). Left inset: TEM image of an irradiated nanoparticles-GFP solution. Right inset: TEM image of irradiated gold nanoparticles in the absence of fluorescent proteins. Scale bars represent 200 nm.

Figure 3

Fig. 3. Proposed model for optical pulse-driven adsorption of fluorescent proteins on gold nanoparticles. (a) Gold nanoparticles coated with PEG-TA within a solution of fluorescent proteins. (b) Irradiation by a single femtosecond pulse results in a cavitation bubble and shock wave formation, leading to partial removal of the PEG molecules. (c) The fluorescent proteins are adsorbed onto the particle and their fluorescence is quenched. (d) Following additional pulses, the quenched proteins are gradually degraded. (e) The addition of DTT to the solution results in protein desorption from the nanoparticles.

References

1. Controlled fabrication of gold nanoparticles and fluorescent proteins conjugates
   Gili Bisker, Limor Minai and Dvir Yelin
   Plasmonics , (2012)

Gold nanoparticles find a wide range of applications in optics and photonics; however, their detailed interaction with intense laser light is only partially understood. Previous works have studied the effect of intense pulse trains on gold nanoparticles at a wide range of illumination parameters, and observed diverse optical and morphological changes. In this work we study, for the first time, the interaction between single femtosecond pulses and gold nanoparticles. Using transmission electron microcopy and optical spectroscopy, we have found that nanoparticles illuminated by 50 fs pulses with fluence of less than 0.15 J/cm2 per pulse (3 TW/cm2) undergo morphological changes which affect their extinction spectra. Experimentation with particles of different diameters show similar qualitative effects, which are more pronounced for larger particles. Pulses at different excitation wavelengths were found to induce different effects for resonance and off-resonance conditions. The presented results provide valuable experimental data on the complex pulse-particle interaction and would be helpful for better understanding of the physical processes that are involved.

Figure 1

To study the effect of femtosecond pulses on the gold nanoparticles, nanoparticle solutione were placed in a 1 mm thick rectangular glass cuvette, which was placed at the focal plane of an optical arrangement (Figure 1). This experimental setup allowed the illumination of the particles with single pulses while observing changes in the extinction spectra. Light from a visible broadband source was coupled into a multimode optical fiber whose distal end was imaged into a small spot on the cuvette using two achromatic lenses L1 (75 mm focal length) and L2 (50 mm focal length). The focal depth was approximately 5 mm, longer than the total cuvette thickness (1 mm). The light transmitted through the cuvette was then collimated (lens L3, 75 mm focal length) and refocused (lens L4, 150 mm focal length) into a multimode fiber, which was used as the input of a commercial spectrophotometer having a maximal capturing rate of 250 spectra per second. The wavelength-dependent extinction coefficients were computed by taking the logarithm of the ratio between the transmitted spectra of pure water and that of the nanoparticle solution.

Figure 2

The extinction spectra of solutions having particle diameters of 16, 23, 45 nm following illumination of single 50 fs pulses of 90 mJ/cm^2 per pulse (1.8 TW/cm^2), at a center wavelength of 550 nm, are shown in parts a, b, and c respectively of Figure 2. A comparison between the extinction spectra before illumination (black line) and immediately after (less than a second) 1 pulse (blue), 2 pulses (green), and 10 pulses (red) reveal noticeable differences in the extinction spectra after each pulse. The effect of the first pulse was the most significant for all three solutions, and included noticeable blue shifts of the spectra, spectral narrowing by 5-25%, decrease in the maximum extinction values by approximately 5%, and a substantial decrease of the extinction at near-infrared wavelengths by 70-90%. The effect of the second pulse was less pronounced but had a similar trend, with the exception of the peak extinction values which slightly increased compared to the peak values after 1 pulse. Similar effects, although gradually decreasing in magnitudes, were observed after each additional pulse, until no changes were visible after more than 10 pulses.

Figure 3

The effect of the resonant illumination at 550 nm could be noticed by comparing the TEM images of the GNP-45 solution before illumination (Figure 3a), after 1 pulse (3b), 2 pulses (3c), and 10 pulses (3d). Before illumination (3a), the solution contained mainly nanospheres, but also a certain quantity (approximately 20%) of different particle shapes, including triangles, rods, and icosahedrons. After the first pulse illumination, a more homogeneous assortment of particles was observed, with a smaller abundance of nanorods and of other particles with sharper corners or edges (3b). This trend continued with each additional pulse (3c,d), until almost all of the particles were transformed into spheres with a narrower size distribution. A similar effect was evident by visual examination of the TEM images of the resonantly illuminated drops from the GNP-23 and GNP-16 solutions (data not shown). Scale bars represent 100 nm.

References

1. The effect of single femtosecond pulses on gold nanoparticles
   Omri Warshavski, Limor Minai, Gili Bisker, Dvir Yelin
   J. Phys. Chem. C 115, 3910 (2011)

Specifically targeting and manipulating living cells is a key challenge in biomedicine and in cancer research in particular. Several studies have shown that nanoparticles irradiated by intense lasers are capable of conveying damage to nearby cells for various therapeutic and biological applications. In this work we utilize ultrashort laser pulses and gold nanospheres for the generation of localized, nanometric disruptions on the membranes of specifically targeted cells. The high structural stability of the nanospheres and the resonance pulse irradiation allow effective means for controlling the induced nanometric effects. The technique is demonstrated by inducing desired death mechanisms in epidermoid carcinoma and Burkitt lymphoma cells, and initiating efficient cell fusion between various cell types. Main advantages of the presented approach include low toxicity, high specificity and high flexibility in the regulation of cell damage and cell fusion, which would allow it to play an important role in various future clinical and scientific applications.

Figure 1

Figure 1: Schematic of the optical system for nano-manipulations of cells. A beam from an optical parametric amplifier (OPA) at wavelength 550 nm (green line) was scanning the sample using two scanning mirrors. A single lens (L1) was used to reduce the beam’s diameter on the sample to 250-350 micron. For two-photon imaging, the beam from the Ti:Sapphire oscillator was picked using a flipped mirror (M), scanned and magnified using lenses L2 and L3, and focused by inserting an objective lens before the sample. The two-photon fluorescence signal was detected by a photo-multiplier tube (PMT) after replacing the high reflectivity (HR) mirror by a dichroic short-pass mirror (LP-700). After irradiation, time lapse imaging was conducted by replacing the dichroic mirror with different filter cubes optimized for the specific fluorescence markers. Schematic illustrations show (I) the irradiation pattern of the pulse beam, (II) the irradiated nanoparticle-targeted cells, and (III) the extent of the nanometric effect induced by each nanoparticle, including the near field enhancement beyond the ionization threshold (red regions), the spherical shock wavefront (gray sphere), and the affected area on the membrane (dashed circle).

Figure 2

Figure 2: Gold nanoparticles on carcinoma cell membranes. Two-photon imaging of A431 cells incubated with (a) anti-EGFR coated gold nanoparticles, (b) no nanoparticles and (c) PEG-coated gold nanoparticles. Scale bars in a-c represent 30 micron. (d) Scanning electron microscopy of a cell membrane targeted by anti-EGFR gold nanoparticles. (e) Scanning electron microscopy using the back scattering detector reveals bright reflections indicative of gold particles (marked by arrows). Scale bars in d-e represent 100 nm.

Figure 3

Figure 3: Damaging carcinoma cells using anti-EGFR coated nanoparticles and resonance femtosecond pulse irradiation. (a) Gold nanoparticle-conjugated cells irradiated by 16 pulses. (b) Non-conjugated cells, 16 pulses. (c) Conjugated cells irradiated by 4 pulses. (d) Non-conjugated cells, 4 pulses. (e) Conjugated cells irradiated by 1 pulse. (f) Non-conjugated cells, 1 pulse. Multi-nucleated cells margins are marked by white dashed curves in (c). Scale bars represent 50 micron. Red nuclei indicate necrotic cells. Green stain indicates apoptosis. Panels a-d show cells 5 hours after irradiation. Panels e, f show cells 23 hours after irradiation.

Figure 4

Figure 6: Induction of cell fusion. (a) Time sequence of fluorescence images of fusing BJAB cells superimposed on phase contrast images, following irradiation by 5 pulses in the presence of non-specific gold nanoparticles. Plasma membranes were green labeled. (b) Formation of a hybridoma cell. Time sequence of fluorescence images of human BJAB cells (red nuclei) and murine (NSO) cells (blue nuclei) superimposed on phase contrast images, following irradiation of 5 pulses in the presence of non-specific gold nanoparticles. Scale bars represent 10 micron.

References

1. Optical Nano-Manipulations of Malignant Cells: Controlled Cell Damage and Fusion
   Limor Minai, Daniella Yeheskely-Hayon, Lior Golan, Gili Bisker, Eldad J. Dann and Dvir Yelin
   Small , (2012)
2. High levels of reactive oxygen species in gold nanoparticle-targeted cancer cells following femtosecond pulse irradiation 
   Limor Minai, Daniella Yeheskely-Hayon and Dvir Yelin
   Scientific Reports 10.1038, srep02146(2013)

Figure 1

The attachment between two different cells via a bispecific nanoparticle is illustrated in Figure 1a. Following the addition of the nanoparticles to the growth medium of an equally (1:1) mixed BJAB (green labeled) and human monocyte-derived dendritic (blue labeled) cell population (Figure 1b), approximately 30% of the cells formed pairs or small clusters of physically attached cells (Figure 1c). Significantly lower cell attachment levels of 7, 6.5, and 4.5% were observed when incubating similar cell mixtures with only single antibody (anti-CD20) nanoparticles, nonspecific antiepidermal growth factor receptor (EGFR)-coated nanoparticles, and in the absence of nanoparticles, respectively.
A closer view of a fi xated BJAB–DC pair is shown in the false-color scanning electron microscopy (SEM) image ( Figure 2a), revealing the light-toned BJAB cell (green) attached to the larger and feature-rich DC (blue). A magnified view (Figure 2b) of the region marked by a rectangle in Figure 2a, overlaid by an atomic-number-sensitive image of the same fi eld of view, reveals a relatively uniform distribution of the individual gold nanoparticles (orange yellow) on the plasma membranes of both cells, without any noticeable particle aggregation. Based on these images, we estimate particle densities of approximately 40 and 12 particles per μm^(−2) on the BJAB and the DCs, respectively. As expected, due to the presence of numerous identical antibodies on each particle, attachment between cells of a similar type after incubation with bispecific nanoparticles was also evident: more than 40% of the BJAB cells and approximately 2% of DCs appeared as part of small (2–10 cells) clusters.

Figure 2

Cells engaged in a physical contact do not normally fuse together in the absence of specific biological or chemical induction signals. Consequently, spontaneous fusion
events were not observed at any of the cell cultures even after several hours. In a previous work we have shown that cells within a homogeneous solution incubated with monospecific (single antibody coating) gold nanospheres were led to fuse following irradiation by a few high-intensity femtosecond pulses that were tuned to the plasmonic resonance wavelength of the nanoparticles. The fusion effi ciency of cells of different types was extremely low in these experiments due to the random and transient physical contacts between the cells.

Figure 3

To demonstrate specific cell fusion, the entire volume of the culture containing the coupled BJAB–DC pairs was washed off the unbound nanoparticles and irradiated by a sequence (ten pulses, 1 kHz repetition rate) of intense (12 mJ/cm^(− 2)), ultrashort (50 fs), resonant (545 nm wavelength) laser pulses.Time-lapse fl uorescence-phase contrast imaging of two exemplary fusion events between BJAB and DCs is shown in Figure 3a. Data analysis based on 1000 cells revealed that fusion between BJAB and DCs was less common (7%) than fusion between homogeneous pairs (or clusters) of BJAB cells only (20%). Moreover, DC–DC fusion was not detected at all, unless a BJAB cell was involved in the fusion process into a DC–DC–BJAB cell hybrid. The irradiated BJAB cells and DCs that were in physical contact showed fi rst signs of morphological changes immediately after irradiation, toward the formation of hybrid cells with a unifi ed cytoplasm after 20–30 min (Figure 3a). A schematic model of the fusion process between engaged cells irradiated by femtosecond pulses is illustrated in Figure 3b, which depicts the local disruption of the cells’ plasma membranes initiated by the several tens of nanometers in diameter cavitation bubble that has been formed around the particle. In addition to DCs, foreign or tumor antigens could also be presented to the adaptive immune system by macrophages, following attachment and fusion with the target cells. To demonstrate the induction of fusion between macrophages and malignant B cells, monocyte-derived macrophages expressing CD86 surface receptors were mixed with CD20-expressing BJAB cells and incubated within a medium containing bispecific anti-CD20–anti-CD86 nanoparticles. After 15 min of incubation, nearly 50% of the macrophages in the culture were found attached to BJAB cells (Figure 4a,b), a significantly higher rate compared with coupling in a nanoparticle-free medium (20%). In contrast, the use of monospecific anti-CD86 nanoparticles resulted in 26% macrophage–BJAB cell attachment and 30% macrophage–macrophage attachment. After irradiation by a single sequence (ten pulses, 1 kHz repetition rate) of femtosecond pulses, approximately 9% of the macrophage–BJAB cell pairs were successfully fused. A typical fusing pair of cells is shown in a time-lapse image sequence in Figure 4c, which reveals the plasma membrane fusion (phase contrast, top row), cytoplasmic (green fluorescence, second row) transfer from the BJAB cell during fusion, and gradual passage of the nuclear blue marker from the macrophage nucleus to the BJAB cell nucleus (third row). Consistently, the fused cells remained viable more than 24 h after irradiation in all fusion experiments.

Figure 4

In summary, we have presented a novel technique for selective fusion between antigen-presenting cells and tumor cells, by using specifi cally designed nanoparticles for the attachment of selected cells, and fusion of their plasma membranes following irradiation by a short sequence of resonant femtosecond laser pulses. The low toxicity of the gold particles, high specificity, efficiency, and relative simplicity of this approach would make it useful for a wide range of biomedical applications, and open new possibilities in biotechnology and in fundamental biological research. This approach would also broaden the use of nanotechnology for various biomedical applications by offering effective means for triggering and controlling desired interactions on a nanometer scale, potentially resulting in a more accurate research procedure and
less invasive medical intervention.

References

1. Optically Induced Cell Fusion Using Bispecific Nanoparticles
   Daniella Yeheskely-Hayon, Limor Minai, Lior Golan, Eldad J. Dann, and Dvir Yelin
   Small 10.1002, 201300696 (2013)

Utilizing Fourier-domain interferometry, spectrally encoded endoscopy (SEE) was shown capable of video-rate three-dimensional imaging, albeit at limited depth of field due to the limited spectral resolution of the detection spectrometer. We show that by using dispersion management at the reference arm of the interferometer, the tilt and curvature of the field of view could be adjusted without modifying the endoscopic probe itself. By controlling the group velocity dispersion, this technique is demonstrated useful for imaging specimen regions which reside outside the system’s depth of field. This approach could be used to improve usability, functionality and image quality of SEE without affecting probe size and flexibility.

Figure 1: Optical setup for controlling field of view by dispersion management

Our benchtop spectrally encoded imaging system (figure 1) consisted of a broadband titanium sapphire oscillator (Femtolasers Rainbow, 300 nm bandwidth, 800 nm center wavelength) coupled to a 50/50 single-mode optical fiber coupler within a Michelson interferometer arrangement. The sample arm consisted of a fiber collimator, a focusing lens and a 1200 lines/mm transmission diffraction grating (G), with 30º Littrow’s angle at 830 nm. The spectral line was scanned in the y axis by a galvanometric scanner which was placed between the grating and the sample. The reference arm consisted of a collimating lens, a variable neutral density (ND) filter, a dispersion control unit, and a mirror mounted on a translation stage. Interferograms of the reflected spectra from the sample and the reference arms were detected by a custom spectrometer, comprised of a collimating lens, an 1800 lines/mm transmission diffraction grating, a focusing lens and a high-speed line camera (Basler Sprint spL4096).

Figure 2: Images of a bovine stapes. Scale bars represent 1 mm. FP – foot plate, C – crus, H – head.

In order to demonstrate effective three-dimensional imaging of a middle ear ossicle without moving the probe, a bovine stapes was attached to a flat surface (figure 2). When an endoscopic probe is inserted into the middle ear through the Eustachian tube, its axis would be nearly perpendicular to the main axis of the stapes which spans along its head, the two crura, and the foot plate. To simulate such configuration, we have imaged the stapes at approximately 90° angle between the probe and the stapes main axis. With no GVD difference between the interferometer arms, the foot plate could be clearly observed (figure 2a), however the two crura, the neck and the head of the stapes, all fall outside the image range and appear dark. With a grating pair separated by 30 mm in the reference arm, the stapes could be fully visualized (figure 2b), while maintaining the same view angle between the probe and the ossicle. The dark horizontal lines on the flat surface below the stapes in figures 2a-b correspond to the reference plane which results with no spectral modulation, and was filtered out by our data processing algorithms.

References

1. Dispersion management for controlling image plane in Fourier-domain spectrally encoded endoscopy
   Michal Merman and Dvir Yelin
   Opt. Express 19, 4777 (2011)

Optical microscopy of blood cells in vivo provides a unique opportunity for clinicians and researchers to visualize the morphology and dynamics of circulating cells, but is usually limited by the imaging speed and by the need for exogenous labeling of the cells. Here we present a label-free approach for in vivo flow cytometry of blood using a compact imaging probe that could be adapted for bedside real-time imaging of patients in clinical settings, and demonstrate subcellular resolution imaging of red and white blood cells flowing in the oral mucosa of a human volunteer. By analyzing the large data sets obtained by the system, valuable blood parameters could be extracted and used for direct, reliable assessment of patient physiology.

Figure 1

Fig. 1. Image acquisition in SEFC. (a) A single line within a blood vessel is imaged with multiple colors of light that encode lateral positions. (b) A single cell crossing the spectral line produces a two-dimensional image with one axis encoded by wavelength and the other by time.

Figure 2

Fig. 2. In vivo noninvasive imaging of blood flow in a single vessel. (a) Raw image acquired during 4 seconds of flow. (b) Crop of the raw data acquired during the period marked by the dashed rectangle in a. (c) Crop of the raw data acquired during the period marked by the dashed rectangle in b. (d) Crop of the raw data marked by the dashed rectangle in c. Inset, an image of the blood vessel as seen through the auxiliary green channel. Dashed red line indicates the transverse location of the spectrally encoded line.

Figure 3

Fig. 3 In vivo imaging in microvessels. (a) RBC flow in a vessel which was slightly deformed by applying external pressure. (b) A histogram showing distribution of RBC diameters. (c) Extracting the fractional area occupied by RBCs in a vessel using manual segmentation for assessing hematocrit levels. Red regions correspond to areas occupied by RBCs. (d) Individual RBCs flowing in a small-diameter capillary.

Figure 4

Fig. 4. In vivo imaging of WBCs. (a) Imaging close to the wall of a post capillary venule allows forming a plot of the total intensity across the image as a function of time. Good correlation between the appearances of WBCs (magnified cell images below the plot) and intensity peaks was obtained. (b) Averaged flux of WBCs for different vessel diameters. (c) Diameter histogram of the imaged WBCs. (d) A single rolling WBC (outlined by a dashed blue line) near a vessel wall, characterized by an elongated shape which indicates its low relative velocity (approximately 0.18 mm/s) in contrast to the nearby fast RBC flow (approximately 1.1 mm/s). (e) Two attached WBCs flowing near a vessel wall. (f) A single WBC flowing in a small capillary with a downstream plasma gap (RBC depleted region).

References

1. Flow cytometry using spectrally encoded confocal microscopy
   Lior Golan and Dvir Yelin
   Opt. Lett. 35, 2218(2010)
2. Noninvasive imaging of flowing blood cells using label-free spectrally encoded flow cytometry
   Lior Golan, Daniella Yeheskely-Hayon, Limor Minai, Eldad J Dann, and Dvir Yelin
   Biomed. Opt. Express , (2012)
3. Measuring blood velocity using correlative spectrally encoded flow cytometry  
   Tal Elhanan  and Dvir Yelin
   Optics Letter 39, 4424  (2014)
4. Reflectance confocal microscopy of red blood
   cells: simulation and experiment 
   Adel Zeidan and Dvir Yelin
   Biomedical Optics Express  6, 4335 (2015)

In its current mode of implementation, SEE has several limiting factors which need to be addressed before its clinical promise could be realized. First, the use of wavelength to encode space imposes some difficulties on wavelength-sensitive imaging modalities. For example, fluorescence spectrally encoded imaging required a sophisticated optical setup for frequency-encoding [19]. Additionally, the use of spatially coherent illumination through a single mode fiber causes pronounced speckle noise, small depth of field, and poor signal collection efficiency which often requires the use of lasers, supercontinuum generation sources, or high power super-luminescent diode arrays. One possible solution for addressing these issues includes the use of a double-clad fiber for spatially coherent sample illumination and incoherent signal collection. While double-clad SEE was demonstrated capable of speckle-free imaging with large depth of field, the endoscopic probe itself suffered from significant cross-talk between the illumination and the collection channels. Back reflections from the probe’s optics, which were efficiently collected by the large area and the high numerical aperture of the inner cladding, resulted with high image noise and required continuous background subtraction during image acquisition.

Figure 1

We show a new approach for SEE, termed multiple-channel SEE (MC-SEE), which addresses many of the image quality concerns of SEE and further expands its functionality. While current forms of SEE involve spectral encoding in both the collection and illumination channels independently, effective encoded imaging is still feasible using only one encoded channel. The concept of imaging with a single encoded channel is schematically illustrated in Fig. 1, showing space-to-wavelength encoding of broadband light (e.g. fluorescence) emanating from a specimen.

Figure 2

MC-SEE images of a portion of a 1 Euro cent coin using spectrally encoded coherent (top panel) and incoherent front illumination (bottom panel) are shown in Fig. 2. Comparison between the two images reveals significant reduction in speckle noise using incoherent illumination, resulting in a more natural appearance with better discrimination of surface texture.

Figure 3

An image of a USAF-1951 fluorescence resolution target was acquired using incoherent excitation light (Nikon Fiber Illuminator, 130 W mercury lamp) and a 360-440 nm bandpass excitation filter, with exposure time of 100 ms per line (Fig. 3a). Along its wavelength (horizontal) axis, the fluorescence image revealed an intensity profile which resembles the emission spectrum of the fluorophore, resulting with a field of view of approximately 1.5 mm in the x axis. To demonstrate the potential of our approach to image weak fluorescence signals of biological specimen, we have imaged a culture of fixed human epithelial cells of breast adenocarcinoma origin (MDA-MB-231), whose membranes were labeled with a green fluorescent marker (DiOC18), using a 465-495 nm bandpass excitation filter. The resulting MC-SEE fluorescence image of the cells, acquired with 500 ms line exposure time, is shown in Fig. 3b, next to an image of the same field of view using a conventional epi-fluorescence microscope (NA = 0.45, Fig. 3c). While the numerical aperture of the SEE lens (NA = 0.25) and the signal-to-noise ratio (6.3 dB) were too low to resolve sub cellular structures, single cells could be easily resolved and a good match was obtained between the two images.

References

1. Multiple-channel spectrally encoded imaging
   Avraham Abramov, Limor Minai, and Dvir Yelin
   Opt. Express 18, 14745(2010)

Endoscopes help medical procedures to be less invasive, thereby reducing the risk of complications as well as costs and recovery times, but their application is limited in part by their size and inflexibility and by their inability to provide a three-dimensional perspective. We study a new type of endoscopy that enables video-rate, three-dimensional images to be transmitted from flexible probes that are comparable in diameter to a human hair. This technology opens up the possibility of moving
operations to an outpatient setting, reducing requirements for anesthesia, and minimizing tissue damage.

The first endoscope was invented almost 50 years ago and consisted of a bundle of optical fibers. Miniature endoscopes still use bundles of optical fibers to transmit a two-dimensional image, but larger endoscopes now employ solid-state, charge-coupled-device cameras for superior image quality. Fiber-bundle endoscopes with sub millimeter diameters have been used for a variety of clinical applications. However, they have not been widely adopted owing to their rigidity and inadequate image quality, which is due in part to relatively low numbers of pixels and superimposition of a honeycomb pattern (a pixelation artefact).

Figure 1

Spectrally encoded endoscopy (SEE) is a miniature-endoscopy technique that overcomes many of the limitations of fiber-optic imaging bundles. With SEE, polychromatic light emanating from a single optical fiber is configured such that each color (wavelength) is projected to a different location on the tissue surface (Fig. 1a). Light reflected from the patient can then be decoded outside the body by using a spectrometer, to form one line of the endoscopic image. The second dimension is obtained by moving the fiber using a mechanical-transduction mechanism, such as an external motor or a galvanometer. The diameter of the SEE probe can be as small as that of the optical fiber, which is typically in the range 80–250 μm. Even though the probe diameter is small, the number of pixels in a SEE image is larger than that from fiber bundles, being dependent only on the spectral width of the light source and the ability of the probe to separate out the components of different wavelengths. Spectral encoding can also provide depth information by using optical interferometry. Photographs of a prototype miniature SEE probe are shown in the insets of Fig. 1a. Light from a single-mode fiber, expanded through a 1.8-mm-long silica spacer, is focused by a gradient-index (GRIN) lens and then diffracted by a transmission grating (1,000 lines per mm) fabricated on the tip of the probe at Littrow’s angle. The maximum diameter of the probe is 350 μm. In combination with a broad-bandwidth (700–900 nm) titanium–sapphire laser, Michelson interferometer and high-speed spectrometer (see supplementary information), this miniature probe can be used to obtain volumetric images with about 400,000 resolvable points at video rates (30 frames per second).

To demonstrate the potential of SEE for minimally invasive applications, we imaged
metastatic ovarian tumour nodules on the peritoneum of a living mouse. The SEE probe was delivered through a modified 23-gauge needle into the abdominal cavity. A three-dimensional surface image of the parietal peritoneal wall, obtained with the SEE probe in vivo, is shown in Fig. 1b: several raised tumor nodules are evident and the inset shows a histological section.

The size and flexibility of this device will allow safer navigation through delicate intraluminal structures such as the fallopian tube, the salivary, mammary and pancreatic ducts, and safer fetal, pediatric and neurosurgical interventions. Navigation mechanisms are still needed that add minimal bulk without compromising flexibility. Color could be introduced by using three separate broad bandwidth sources, each centered at visible red, green or blue wavelengths. SEE using fluorescent light should also be useful for imaging labeled molecular species, reporters and molecular beacons.

References

1. Three-dimensional miniature endoscopy
   D. Yelin, I. Rizvi, W. M. White, J. T. Motz, T. Hasan, B. E. Bouma, and G. J. Tearney
   Nature 443, 765(2006)
2. Miniature forward-viewing spectrally encoded endoscopic probe
   Adel Zeidan  and Dvir Yelin
   Optics Letter 39, 4871 (2014)