DEVELOPMENT OF A SINGLE-MODE
INTERSTITIAL ROTARY PROBE FOR IN VIVO
DEEP BRAIN FLUORESCENCE IMAGING
Mémoire présenté
à la Faculté des études supérieures et postdoctorales de l’Université Laval dans le cadre du programme de maîtrise en biophotonique
pour l’obtention du grade de Maître ès sciences (M.Sc.)
FACULTÉ DES SCIENCES ET DE GÉNIE UNIVERSITÉ LAVAL
QUÉBEC
2013
c
Ce mémoire rend compte de l’expertise développée par l’auteur au Centre de recherche de l’Institut universitaire en santé mentale de Québec (CRIUSMQ) en endoscopie fi-brée. Il décrit la construction d’un nouveau type de microscope optique, le Microscope Interstitiel Panoramique (PIM). Par la juxtaposition d’un court morceau de fibre à gra-dient d’indice et d’un prisme à l’extrémité d’une fibre monomode, la lumière laser est focalisée sur le côté de la sonde. Pour former une image, cette dernière est rapidement tournée autour de son axe pendant qu’elle est tirée verticalement par un actuateur piézo-électrique. Ce design de système rotatif d’imagerie interstitielle peu invasif est un effort pour limiter les dégâts causés par la sonde tout en imageant la plus grande région possible en imagerie optique cérébrale profonde.
This thesis documents the expertise developed by the author at the Centre de recherche de l’Institut universitaire en santé mentale de Québec(CRIUSMQ) in fibered endoscopy, particularly the design and construction of a new kind of optical microscope: The Panoramic Interstitial Microscope (PIM). Through the juxtaposition of a short piece of Graded-Index fibre and a prism at the end of a single-mode fibre, laser light is focussed on the side of the probe. To form an image, the latter is quickly spun around its axis while it is being pulled vertically by a piezoelectric actuator. This minimally invasive fluorescence rotary interstitial imaging system is an endeavor to limit the damage caused by the probe while imaging enough tissue to provide good context to the user in deep brain optical imaging.
This three year journey in biophotonics offered me the opportunity to grasp a fragment of the fundamental research milieu. I had the chance to encounter deeply perfervid individuals, especially my supervisors Daniel Côté and Yves De Koninck. My gratitude goes to all of them, along with my wish to see their efforts bear fruit.
The author was supported by a post-graduate award from Le Fonds de recherche du Québec - Nature et technologies (FRQNT).
1 Introduction 1
1.1 Fluorescence Imaging in Biology . . . 2
1.1.1 Principles of Fluorescence . . . 2
1.1.2 Wide-field Microscopy . . . 4
1.1.3 Confocal Microscopy . . . 6
1.2 Fluorescence Imaging of the Live Nervous System . . . 8
1.2.1 Surgical Access . . . 8
1.2.2 Miniature Lenses . . . 9
1.2.3 Fibered Systems . . . 10
1.3 Non Imaging Fibered Probes . . . 13
2 Probe Design and Manufacture 15 2.1 Modelling . . . 16
2.1.1 Illumination Beam Profile . . . 18
2.1.2 Optimal Working Distance . . . 21
2.1.3 Illumination and Collection in Tissue . . . 23
2.2 Probe Assembly . . . 25 2.2.1 Fusion Splicing . . . 25 2.2.2 Angle Cleaving . . . 27 2.2.3 Optical Testing . . . 28 2.2.4 Coating . . . 30 2.2.5 Connector Installation . . . 30
3 Probe Experimental Characterization 31 3.1 Working Distance in Tissue . . . 31
3.2 Resolution . . . 34
3.3 Model Validation . . . 35
4 Panoramic Interstitial Microscope 37 4.1 Excitation Light Delivery and Collection . . . 37
4.1.1 Fibre-Optic Rotary Joint . . . 39
4.2.1 Rotation . . . 40
4.2.2 Axial Motion . . . 41
4.2.3 Probe Guide . . . 42
4.3 Image Generation . . . 42
4.3.1 User-Defined Parameters Input . . . 44
4.3.2 Data Acquisition . . . 44
4.3.3 Reconstruction . . . 45
5 System Integration and Imaging 46 5.1 Signal Detection and Reconstruction . . . 46
5.2 Fluorescent Beads Imaging . . . 47
5.2.1 Resolution Artifacts . . . 49 5.2.2 Reconstruction Jitter . . . 49 5.2.3 Probe/Sample Contact . . . 50 5.2.4 Signal Collection . . . 50 6 Conclusion 52 Bibliography 54
A Matlab Model Functions 61
B Fusion Splicing Parameters on Vytran FFS-2000 75
C Probe Production Guide 77
1.1 Principles of Fluorescence . . . 3
1.2 Principle of Wide-field Microscopy . . . 4
1.3 Rayleigh Criterion . . . 5
1.4 Principle of Confocal Microscopy . . . 6
1.5 Available Fibre Types For Probe Design . . . 10
1.6 Scanning Configurations . . . 12
1.7 Common Fibre Tip Configurations . . . 13
2.1 Fibre Refractive Index Profiles . . . 17
2.2 Beam Diameter Along Propagation Axis . . . 19
2.3 Single-Mode Fibre Output Beam Divergence . . . 20
2.4 Illumination Trends . . . 22
2.5 Modelling Intensity Maps . . . 23
2.6 Probe Assembly Schematics . . . 26
2.7 Optical Element Lengths and Number of Steps Interdependence . . . . 27
2.8 Angle-Cleaved Probes Under a Scanning Electron Microscope . . . 28
2.9 Far-Field and Side-View Probe Imaging . . . 29
3.1 Relation Among Working Distance and Element Lengths . . . 33
3.2 Relation Between Lateral Resolution and Element Lengths . . . 34
3.3 Modelled Against Measured Working Distance and Resolution . . . 35
4.1 Panoramic Interstitial Microscope Optical Path . . . 38
4.2 Fiber-Optic Rotary Joint Insertion Loss Ripple . . . 40
4.3 Panoramic Interstitial Microscope Prototypes 3D Models . . . 41
4.4 LabVIEW Software Acquisition Flowchart . . . 43
4.5 Acquisition Signal Timing . . . 45
5.1 Signal Acquisition and Reconstruction Tests . . . 47
Introduction
Optical microscopy is a technique of choice in neurobiology[38]. Its resolution is suf-ficient to observe cellular structures and the photon energy in the visible and near infrared spectrum is compatible with live samples, unlike electron and scanning probe microscopies. In particular, fluorescence imaging is the most versatile contrast tech-nique because of its multi-stage selectivity originating from fluorophore properties and placement. Fluorescence emission can be restricted to an organ, such as dye injection in the vasculature, or to a specific cell type, through genetically-modified mice strains producing fluorescent proteins for example. With fluorescent contrast agents, neurons don’t need to be physically isolated from the surrounding matrix to be detected. Their form, function and network interactions can be investigated in vivo within the intact tissue or even the whole organism often with subcellular spatial resolution, on a mil-liseconds timescale and for months. However, living tissues are highly scattering and light has to be brought within micrometers from the imaging site in order to benefit from the advantages of fluorescence. Current endoscopic strategies are inadequate to study deep brain structures such as the striatum while preserving enough integrity to the organ for the animal to survive.
The Project
This master’s project goal is to build an apparatus enabling deep brain structures imag-ing over days. It involved the design, manufacture and implementation of a minimally invasive fluorescence interstitial imaging system, capable of transporting excitation and signal light through a single optical fibre and creating high resolution images in a matter of seconds. The Panoramic Interstitial Microscope (PIM) is intended for longitudinal
imaging of the animal brain’s deep structures. I designed the mechanical mounting of the microscope, developed the technique to make probes and programmed the acquisi-tion software.
The complexity of the project required to establish collaborations with several ex-perts. Dr. Brett Bouma, professor at Harvard Medical School, holds an unrivaled expertise in scanned systems with cylindrical symmetry. The neurosurgeon Leo Cantin oversaw design constraints to ensure that the endoscope would meets the needs of an-imal neurobiology and be compatible with clinical translation. Collaborations with Professor Réal Vallée, director of the Center for Optics, Photonics and Laser in Que-bec City (microoptical probe fabrication) and with Professor Jean Ruel, director of the Bureau de design (design and assembly of mechanical prototypes) were also valuable. In vivo tests were carried out in collaboration with Dr. Armen Saghatelyan.
1.1
Fluorescence Imaging in Biology
1.1.1
Principles of Fluorescence
Radiation of light by a non incandescent body is termed luminescence. This phe-nomenon is always triggered by an external energy source that "excites" molecules, meaning that they leave a stable ground-state S0 to reach an higher energy level, as
illustrated in Figure 1.1. The input energy source can be electricity (electrolumines-cence), a chemical reaction (chemiluminescence or biolumines(electrolumines-cence), radioactivity (ra-dioluminescence) or another source of light (photoluminescence). Photoluminescence can either be fluorescence or phosphorescence. The latter is a slow process because it implies transition to a triplet energy state. Fluorescence on the other hand is nearly instantaneous, the delay between photon absorption and emission being in the nanosec-ond timescale. When a fluorescent molecule absorbs a photon at a certain wavelength, it emits another photon at a lower energy or frequency. Part of the incident photon energy is lost through vibrational relaxation of the valence electron to the fundamental excited state S1 in the process. For a given fluorophore, the difference in frequency between the
absorption maximum (the excitation wavelength) and the fluorescence emission peak (emission wavelength) is the Stokes shift.
When the molecule is in the excited state, there is some likelihood that it participates in chemical reactions, known as photochemical reactions, particularly with oxygen to form free radicals. The fluorochrome may then lose its fluorescence properties. At any
S0 S1 S2 T1 01 2 3 4 01 2 3 4 01 2 3 4 01 2 3 4 Quenching Intersystem crossing Absorption
Vibrational relaxation and Internal conversion Fluor escence Phosphor escence Non-radiative relaxation Normalized Intensity Wavelength (nm) Excitation Light
Absorption FluorescenceEmission
Stokes Shift
a) b)
Figure 1.1: Principles of Fluorescence. a) Jablonski diagram of different phenomena induced by light absorption, with an emphasis on fluorescence. S0, S1 and S2 are singlet
energy levels and T1 is the representation of a triplet energy level. b) Alexa Fluor 594
dye excitation and emission spectrum.
moment during excitation, a certain proportion of molecules cease to fluoresce, therefore decreasing the signal intensity. This is referred to as photobleaching. Photochemical reactions may also be phototoxic, that is they damage the specimen.
Fast and repeatable absorption-emission cycles, enabling detection sensitivity down to a single particle, partly explains the interest researchers have accorded to fluorescence since its formal discovery in the nineteenth century. Specificity is the other reason. In the 1930s, the manipulation of fluorochromes emerged as a tool in biological investi-gations to stain tissue components such as bacteria. Fluorescent dyes have distinctive excitation and emission spectrum as well as a characteristic time taken to convert an incident photon into a fluorescence one, namely the fluorescence lifetime. The second aspect of specificity is the ability for a fluorophore to be colocalized with molecular compounds.
Labeling Strategies
Several types of fluorescent labels are used in biology. Endogenous molecules are al-ready in the specimen and are inherently specific, but their brightness is generally low. Organic molecules are simple, have a large efficiency and their chemical composition is often optimized. Fluorescent proteins can be genetically encoded to be produced by the organism itself. Quantum dots are nanocrystals made from semiconducting nano-material layers. Their fluorescence efficiency is huge, but they are relatively big and
Figure 1.2: Principle of wide-field microscopy. The light sources represent two inde-pendent configurations. Kohler illumination, required for homogenous lighting at the sample and maximal resolution, is not shown.
their toxicity is a concern, especially for in vivo studies[35].
An exogenous marker has to be implanted in the specimen prior to imaging. It can be injected, like quantum dots in the vasculature. Some molecules have a particular affinity for an organelle (DAPI binds to A-T rich regions of the DNA, Rhodamine 123 to mitochondrion membranes and Nile red to lipids). They can be introduced in cells, or transfected, by electroporation and other similar procedures that create a temporary physical passage through the membrane without killing the cell. Organic molecules can be attached to an antibody directed against the object of interest. Finally, fluorescent proteins can be included in the genome and fused to the protein of interest, or contained in a viral vector. After the body is infected by the virus, the protein of interest is expressed. Once the labeling step successful, samples are brought under a microscope for imaging.
1.1.2
Wide-field Microscopy
Pioneering fluorescence imaging experiments were conducted on a conventional or wide-field microscope. Today’s microscopes illuminate the sample plane uniformly over the whole field of view with an incoherent white light usually generated by a xenon arc lamp or mercury-vapour lamp. The reflected or transmitted image of the specimen is projected on the user’s retina or a CCD camera by at least one lens, as sketched in Figure1.2. Since the illumination is homogenous, the resolution is solely related to the collection lens and detector.
0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 !10 !8 !6 !4 !2 0 2 4 6 8 10 0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 !10 !8 !6 !4 !2 0 2 4 6 8 10 0 0.1 0.2 0.3 0.4 0.5 0.6 0.7 0.8 0.9 1.0 1.1 1.2 !10 !8 !6 !4 !2 0 2 4 6 8 10 26.5%
Uresolved Rayleigh Criterion Resolved
Detected Intensity
Point Sources
Figure 1.3: Illustration of the Rayleigh criterion. As two point sources are brought further apart, their point-spread function (blue) stop to overlap and they can be dis-tinguished in the total signal (dashed red).
Lateral resolution in optical microscopy is defined as the ability to resolve details of the imaged object. It is often calculated by measuring the system response, also called the point-spread function (PSF), to a infinitesimal object. When no factors such as lens aberrations, beam deviation, saturation or under sampling alters the imaging capacity more than optical diffraction, the microscope is said to be diffraction-limited. The resolution of a diffraction-limited microscope is approximated with the Rayleigh criterion
Lateral Resolutionwidefield= 0.61
λ0
NA. (1.1)
where λ0 is the signal wavelength in vacuum and NA the collection objective numerical
aperture. The Rayleigh criterion uses paraxial approximation and assumes incoher-ent irradiation. It describes the resolution as the minimal distinguishable separation between two equally bright spots. At this distance, the detector response will show a signal drop of approximately 26.5% (Figure 1.3). For fluorescence emission in wa-ter at 500nm detected with a 1.0NA objective, the diffraction-limited resolution is 230nm. Because single fluorophores can cycle through many excitation-emission se-quences, ingenious strategies have brought fluorescence microscopy below this funda-mental limit[7,41, 42, 30, 59].
In the wide-field configuration, signal (reflectance, transmittance or fluorescence) is emitted from the entire field of view, both above and below the objective lens focal plane. Light coming from out-of-focus regions reach the detector and greatly blur the image. Signal coming from the focal plane can also diffract in the sample and contribute to reduce the image contrast. For these reasons, it is difficult to precisely define. A few microns thick sample at most is needed to restrict the observation to the focal plane. In transmission configuration, a thick sample also attenuates the overall signal because of absorption and reflexion. Magnification is achieved by series of lenses, not illustrated in Figure 1.2.
Lens Dichroic Mirror
Light source Pinhole
Pinhole Detector
Sample
Figure 1.4: Principle of confocal microscopy. The illumination is reduced to a single point at the focal plane. Signal coming from other planes is rejected by the confocal pinhole in front of the detector.
Fluorescence imaging mode on a wide-field microscope relies on careful selection of the illumination and detection wavelengths to match the fluorophore’s characteristics. A dichroic mirror that reflects the excitation light and transmits the emitted light (or vice-versa) separates both beams. The reflected or transmitted light is completely blocked by the emission filter, and only photons created by the fluorescence process in the sample will reach the camera. Fluorescence intensity is orders of magnitude weaker than transmitted or reflected light and thus sensitive array detectors such as CCD cameras and long exposition times are used. Fluorescence generation is an isotropic process. Photons originally directed away from the detector can scatter and come back towards it.
1.1.3
Confocal Microscopy
Having to cut thin brain slices using a microtome prior to imaging neurons was unac-ceptable by the inventor of the confocal microscope, Marvin Minsky. He modified the conventional wide-field microscope to eliminate blurring from thick samples by improv-ing the axial resolution, or the sectionimprov-ing. He did so by focussimprov-ing the excitation light onto one point of the sample and by detecting the signal light through a pinhole and a single detector. The pinhole is placed in the image plane conjugated to the microscope objective focal plane to reject scattered and out-of-focus signal photons (Fig. 1.4). By scanning the sample under the microscope, he could reconstruct an image. This tech-nique was slow and high intensity light sources for fluorescence excitation were lacking.
The combination of three technological advances in the 80’s democratized the fluores-cence confocal microscope: The advent of high-power laser sources corresponding to excitation wavelengths of practical fluorophores, developments in digital image acquisi-tion and galvanometric mirrors enabling fast beam scanning on the sample[72]. Today, a laser is mechanically scanned in what is called a Laser Scanning Confocal Microscope (LSCM)[55]. Its use is now common in many areas, particularly in biology because of its compatibility with fluorescence.
Contrary to wide-field microscopy, the confocal microscope illumination is not ho-mogenous but rather confined to solid cones on each side of the focal spot. When this illumination pattern is taken into account along with the detection path, the confocal lateral resolution is found to be[71]
Lateral Resolutionconfocal = 0.56
λ0
NA. (1.2)
This modest improvement in lateral resolution is nevertheless insignificant compared to the gain in the depth of field. Instead of needing to mechanically cut specimens, the confocal microscope can acquire optically thin sections that enable 3D reconstruction of images from multiple planes.
To enhance sectioning, another strategy than rejecting out-of-focus background is to not generate any in the first place. Planar illumination techniques such as light-sheet microscopy[52], where the excitation light is launched perpendicular to the detection optics, is one example. However, the most famous example of this strategy is Two-Photon Excitation Fluorescence (TPEF), discussed next.
Multiphoton Fluorescence
Absorption is usually achieved one photon at a time. The resulting fluorescence is a linear process, scaling with the excitation intensity and the fluorophore conver-sion efficiency. Interestingly, when the photon density is high enough (in the or-der of109Watts/cm2), two photons can be absorbed simultaneously and their energy
summed up, hence the "Two-photon" appellation. [19] TPEF scales with the intensity squared. It can occur only if two photons arrive in the vicinity of the molecule in the order of the femtosecond, the time it takes for absorption to happen. Because of its low probability, TPEF is circumscribed to the focal spot. This reduces photobleaching as out-of-focus planes are not excited. TPEF also diminishes photodamage because for a given fluorophore, each photon is almost twice as less energetic as the single-photon excitation equivalent. The confinement of the excited volume has a third advantage: it
eliminates the need for a pinhole. Fluorescence photons originate from the focal spot only. The overall intensity is lower in TPEF, but every photon counts even if it scatters. The detector can be placed right next to the microscope objective to collect as much light as possible. The major drawback of multiphoton fluorescence is that expensive pulsed laser light sources are needed to achieve such a high photon density. Pulsed light is also altered when traveling in standard optical fibres.
1.2
Fluorescence Imaging of the
Live Nervous System
In vivomicroscopy is an stirring tool for neurological research because it can distinguish individual cells. In a living sample, it can do so fast enough to observe dynamic events in the millisecond timescale. If carried out over several days, relationships between events and disease unfolding can be observed. These inherent advantages of fluorescence microscopy drive the development of strategies to mitigate the major obstacle related to its use, the short penetration depth.
To image the live CNS in a transparent animal such as the zebrafish is a com-pelling tactic[77,24]. Unfortunately, most animals are opaque to visible light, especially mice and rats where disease animal models and genetically modified strains are readily available. One can alternatively rend the specimen transparent by a pigment clearing procedure[21, 16] to allow profound optical imaging, but this technique is only com-patible with fixed tissue. The remaining options to work intravitally are to bring light directly where the site of interest is, or to bring the site of interest to light. An elegant solution is to use an endoscope to guide light[73]. When inserted in natural orifices, endoscopes induce little or no damage. However, the brain has no opening and tissue damage scales with the endoscope size. In such cases, the apparatus is re-baptized an interstitial probe.
1.2.1
Surgical Access
A convenient sample to study the brain with optical microscopy is to slice it in thin layers. Every part of the brain can be observed this way and cells can be kept alive for up to eight hours by bathing the slice in an oxygenated cerebrospinal fluid solution. This however destroys network connections between brain regions and prevents longitudinal imaging.
A way to access the animal nervous system without literally cutting it up is to expose a part of it to a regular microscope objective through invasive, life threatening and often terminal surgeries. In the brain, chronic imaging studies can be performed on the cortex by implanting a cranial window[2, 15] or by thinning the skull to a 10-30µm thickness[51]. Deeper cortical regions can be reached by the insertion of a glass prism[13]. Besides, the latter geometry lets multiple cortical layers appear on the same field of view. To image the spinal cord, a laminectomy is the standard procedure. This operation involves exposing the spine and removing one or more of its segments[5]. Recently, Farrar et al. developed a spinal window[23]. With this new tool, the operation is no longer exclusively terminal since the window, essentially a glass coverslip, seals the exposed site and prevents infection. Still, imaging is limited to the posterior surface of the spinal cord.
1.2.2
Miniature Lenses
The surgery footprint on rats and mice can be reduced through the use of a smaller microscope objective. Olympus sells the MicroProbe, a range of miniature micro-scope objectives[10], and GRINTech sells a 0.8 numerical aperture gradient-index lenses assembly[4] to this intent. In the animal brain, it is possible to insert a micro objective and observe a sub-surface region over months. Barretto et al. tracked CA1 hippocam-pal pyramidal neuron dendrites by installing 1.8-mm wide glass guide tubes windows[3]. Kim et al.[40] demonstrated more profound brain imaging with a 1.25-mm side-viewing GRIN rod microendoscope, although not longitudinally. Furthermore, we showed that such a microlens could reduce the laminectomy invasiveness by eliminating the need for bone removal[6]. This method is nevertheless limited to the surface since the microlens diameter, 1.4mm, is roughly equal to the rat spinal cord size.
An implanted microlens captures a tiny and fixed part of the tissue. Even a 1-mm graded-index lens assembly has only a 100-µm diameter field of view and its numerical aperture decreases radially[73]. In addition, a typical microlens is over 350µm in size. Chromatic aberrations are a greater concern because of the glass quantity light has to pass through. Autofluorescence due to mechanical damage and immune response suffered by the issue in the probe vincity[47, 48] is yet another. Despite challenging op-tical limitations, two-photon imaging in miniature lenses has a decade long history[36]. Recent progress enabled us to demonstrate another nonlinear imaging modality called coherent anti-Stokes Raman scattering in a GRIN-based microendoscope.
Figure 1.5: Available fibre types for probe design.
1.2.3
Fibered Systems
Fibered endoscopic strategies are selected because they facilitate clinical translation by providing flexibility and potential miniaturization. Optical fibres have traditionally been used in a bundle as flexible relay tubes for endoscopes. As optomechanical tech-nologies are being refined and new fibre types emerge such as double-clad and photonic crystal fibres, a variety of fibered probe configurations arose for fluorescence excitation, sensing and imaging.
The available fibre types for fluorescence probe development are schematized in Figure 1.5. Single-mode fibres are common in confocal imaging, but not in point-measurement applications owing to their low collection efficiency. Multimode fibres, being step-index or gradient index, enhance the detection intensity but reduce section-ing. Double-clad fibres, with their excitation central core and outer detection core are a good compromise. Graded-index fibres can be cut in short pieces and used as small GRIN lenses. Finally, photonic crystal fibres have been developed to alleviate high intensity pulses distortion for nonlinear imaging. They can adopt numerous structures and reducing them to a single pictogram is an outrageous simplification, justified on the grounds that this work will not dig further into nonlinear excitation.
Fibre Bundles
A single fibre bundle can contain several tens of thousands of 2-4µm diameter fibre cores[32]. Each fibre, made of a guiding core enclosed in a cladding of a lower refractive
index, transfers a point of an image plane at one end of the bundle to the other. This process reconstitutes a pixellated image with dead space corresponding to the cladding regions and has a limited resolution. Besides, signal leaking from a fibre to adjacent ones affects the resolution. On account of those drawbacks, solid-state video cameras are often preferred[18], with a fibre bundle as the illumination source. The one application where fibre bundles still have a prominent place is miniature confocal fluorescence microscopy, firstly demonstrated in 1993[29]: When a focussed laser spot is injected sequentially in each fibre at the proximal bundle surface by scanning the beam, each fiber then acts as a confocal pinhole. Such a microscope can operate with the distal end of the bundle in direct contact with the imaged tissue or by appending lenses to refocus the image of the bundle tip onto the specimen. The latter configuration enables the addition of a 90˚prism to build a side-looking endoscope. Image deconvolution is a frequent tactic employed to improve fibre bundle image quality. Santos et al.[60] recently demonstrated an alternative strategy, HiLo bundle imaging, for out-of-focus background rejection. Bundle imagers as small as 300µm have been used in vivo to visualize many tissues, including blood vessels[44], ovaries[68, 76], lungs[45] and brain[33].
Single Fibres
Another concept for confocal fluorescence microendoscopic imaging is to carry the il-lumination and detection light through a single optical fibre, and to scan the beam distally. The most frequent way a fibre is physically scanned is by a piezoelectric resonant scanner providing a lissajous-patterned motion, as illustrated in Figure 1.6. Helicoidal rotational scanning is another pattern used to be more fast and energy ef-ficient than raster scanning. Alternatively, one can use tiny microelectromechanical (MEMS) mirrors to steer the light beam coming out of a single fibre[43, 1, 64]. The resolution and the field of view are in general better than in fibre bundles, but the distal end, containing the scanning mechanism, is rigid and larger. As previously mentioned, probe size is of paramount importance for CNS imaging. No in vivo demonstration of interstitial distally scanning fluorescence probes in deep brain regions been reported yet.
In the quest of the most noninvasive CNS imaging apparatus, the Schnitzer labora-tory at Stanford University have worked on miniaturizing the confocal microscope for freely moving live animal imaging[31, 25, 26, 56, 28]. Their prototypes are miniature fluorescence microscopes fixed over a cranial window or an implanted GRIN lens on a freely moving animal. Various iterations of their apparatus included fibre bundles, multimode and monomode fibres as well as piezoelectric and MEMS distal scanning. Detection optics and and sensor have also been integrated distally. They even moved
Figure 1.6: Scanning configurations. The focal spot can be scanned before the en-doscopic apparatus or after. The bottom-right panel illustrates two common distal scanning patterns.
Figure 1.7: Common Fibre Tip Configurations the light source on the animal head in their most recent version.
1.3
Non Imaging Fibered Probes
Up to here, focus has been kept on imaging systems. However, quantitative information related to the molecular nature of samples can be extracted from spatially unscanned fluorescence signals. The lack of contextual setting is counterbalanced by simpler and smaller devices. Several groups have developed dozens of fibered and miniaturized optical biomedical devices using linear interaction modalities over the last decade. Some have been built and used, other have only been designed and their capacities modelled. Non imaging fibre probes are most often designed for skin and epithelial tissue. Flat-tipped fibres (Fig. 1.7) are widely used for their simplicity[53]. Each fibre usually delivers or collects light, but some have both duties. Separate delivery and collection fibres make for simpler experimental setups and permit diffuse reflectance or fluores-cence spectroscopy. As the delivery and collection fibres are brought further apart, the signal component from scattered photons increases and deeper portions of the tissue is sampled[39]. Adding collection fibres provides more signal and more information on the sampled volume[54, 67, 8]. Bifurcated geometries and fibre rotation are also strategies used to to enhance depth specificity[11, 34, 53, 66]. Another approach to restrain the interrogation volume is to add a ball-lens at the probe tip [14, 61, 22]. When a single fibre is used, the core and inner clad (in the case of double-clad fibres) sizes are the most significant parameters for sampling depth effects. Finally, diffusing tips are only useful for non sensing interstitial delivery systems for light-based treatments such as
photodynamic therapy (PDT).
In the brain, most fibres are in fact non sensing live animal optogenetic stimulators. A recent breakthrough, achieved by Lechasseur et al., combines an electrode and fluo-rescence detection ability in a dual-core 10-µm tipped fibre[46]. This device paves the way for optogenetic coupling to fluorescence signaling feed-back.
Probe Design and Manufacture
Central in PIM development is the design and manufacture of side-viewing single-mode fluorescence microoptical probes. Most other fluorescence imaging probes being devel-oped for neuroscience are currently based on fibre bundles or on graded-index lenses. By geometry, their field of view scales with the probe area, which itself determines damage. Side-viewing probes offer several advantages in the nervous system such as better contextual setting due to the brain’s layered structure and their possibility of imaging repeatedly the same tissue without destroying it. Most importantly, their sizes can be scaled down without loss in resolution or in field of view. The probe presented in this chapter is 125 µm wide and is covered by a protective needle, bringing the outer diameter between 300 µm and 500 µm. Future probes are expected to be as thin as 200 µm.
Our probe design is inspired from the pioneering work of Emkey et al.[20] and implemented by Tearney et al.[69] and Reedet al.[57] using an interferometric modality (Optical Coherence Tomography). The fibered microoptical assembly is single-mode and therefore non-imaging. The image is reconstructed by fibre rotation and pull-back as described in Chapter 4. Bird et al.[9] provided an early demonstration of fluorescence detection in this configuration. In 2007, Mao et al.[50] inserted a spacer element between the single-mode guiding fibre and the graded-index fibre to increase the working distance.
This chapter describes how the production parameters were estimated with mod-elling and how our model is being refined to study the effects of future design modifi-cations. The second section reports on probes fabrication process.
2.1
Modelling
The purpose of this modelling was to understand what factors are important in probe production in order to obtain a probe having 1-a tight focal spot and 2-the optimal working distance in tissue. The ray transfer analysis method provides accurate be-havioral trends and is simple to implement. Paired with experimental verification, it easily fulfilled our needs. Beam illumination propagation and fluorescence collection modelling have been conducted using Matlab (Mathworks). Important functions of the code are included in Appendix A. The vast majority of algorithms are based on ray transfer matrix analysis for Gaussian beams. The ray transfer technique implies paraxial approximations sin θ ≈ θ, tan θ ≈ θ and cos θ ≈ 0, which holds true in our case since the beam never experiences focussing angles greater than 10◦. The Gaussian
approximation is however not appropriate, but compensated (see section 2.1.1).
The ray-tracing technique makes use of two perpendicular planes to the optical axis of the system. A ray is defined by a distance x from the optical axis and an angle θ. The ray is transferred from the input plane (x1, θ1) to the output plane (x2, θ2) by an
optical system characterized by a 4-element transfer matrix, often referred to as the ABCD matrix " x2 θ2 # = " A B C D # " x1 θ1 # . (2.1)
Transfer matrices are known for all linear optical elements[65]. If the ray is a Gaussian beam, a complex beam parameter q can be defined and propagated in the system:
" q2 1 # = k " A B C D # " q1 1 # . (2.2)
The complex beam parameter q is 1 q = 1 R − iλ0 πnω2, (2.3)
where R is the radius of curvature of the gaussian beam at that position, λ0 is its
wavelength in vacuum, n the refractive index of the material and ω its waist size; where the beam intensity drops to 1/e2 of the central axis peak value. Equation 2.3 is mostly
used in our model to define the waist size at the focal spot and in the single-mode guide, ω0 where the real part of q vanishes. The transfer matrices are defined for homogenous
refractive index and graded-index media:
Dielectric transfer matrix: "
1 L/n
0 1
#
a) b) c) Fit function : 1.48386 - 2.8707!10"5x2 Standard deviation : 0.0005572 Index of r efraction 1.455 1.460 1.465 1.470 1.475 1.480 1.485
Distance from center (µm)
"35 "30 "25 "20 "15 "10 "5 0 5 10 15 20 25 30 35
Index vs Position in core
Quadratic fit of Index vs Position in core d)
Figure 2.1: Fibre refractive index profiles (vertical axis) of a) single-mode, b) coreless, and c) gradex-index fibres. d) Fit of the graded-index fibre transfer matrix γ factor.
Graded-index fibre transfer matrix: "
cos(γL) sin(γL)/n0γ
−sin(γL)n0γ cos(γL)
#
. (2.5) Mathematical equations follow Siegman’s notation[65]. L is the element length and n its refractive index. The meaning of γ is discussed in section2.1.1.
Three types of fibres are necessary to produce a microoptical probe. We acquired single-mode fibres from Nufern (S630-HP) and Fibercore (SM600), coreless or multi-mode fibre from Nufern (FUD-3582) and graded-index fibre form Corning (InfiniCor 300). The refractive index profile of each fibre type has been measured on an EXFO NR-9200HR optical fibre analyzer (2.1) at 657.6nm. Refractive indices in table2.1 have been extracted from this experiment.
Table 2.1: Fibre parameters for ray-transfer matrix analysis.
Parameter Symbol Units Measured Value
Single-Mode Fibre
Core Mode-field diameter MFD µm 4.2±0.5 µm @ 630nm
Core Refractive index ncore 1.456
Cladding Refractive index nclad 1.450
Multimode/No Core Fibre
Core Diameter Dsp µm 125
Core Refractive index n 1.456
Element 1 Length Lsp1 µm 500 − 550
Element 2 Length Lsp2 µm 125
Graded-Index Fibre
Core Diameter Dgrin µm 62.5
Central Refractive Index n0 1.484
Parabolic Variation Factor n2 2.871 · 10−5
Gamma γ /mm 6.22
GRIN Length Lgrin µm 180 − 220
2.1.1
Illumination Beam Profile
Light is carried to and from the probe microoptical components by a single-mode fibre (SMF). The first microoptical element is a coreless or No Core fibre (NCF), which is in fact a 125-µm core multimode fibre (see Figure 2.1 b)). It is sometimes referred to as a "spacer" too. The next element is the lens, made from a short graded-index fibre (GRIN) piece. A second spacer is affixed to serve as a deflecting prism. This optical system images the extremity of the SMF in the sample. Figure 2.2 shows how the illumination beam is propagated through the optical system.
The first spacer length (50–550 µm) is set to allow maximal beam expansion without excessive leakage in the GRIN cladding, which is 62.5 µm in this case. The second spacer is as short as possible (125 µm) since it acts only as a prism. GRIN length (180–220 µm) is optimized to place the focal point close to 80 µm into the sample to avoid imaging perturbed tissue while maintaining maximum signal level: the farther the focal point is, the lower is the NA and the more effect absorption and scattering have on imaging. The coreless element for beam expansion is essential to increase the probe working distance[50]. Its effect depends on the input beam parameters. In theory, the complex
B eam diameter (µm) 0 20 40 60 Distance in probe (µm) −200 0 600 750 880 1300
SMF Spacer GRIN Spacer Tissue
Focal spot
Figure 2.2: Beam diameter along propagation axis as modelled by the ray-tranfser matrix method. Distances in probe vary and are only shown for illustrative purposes. beam parameter can be derived from equation 2.3 at the SMF end face because 1
R = 0.
The geometrical core size is not equivalent to the beam size in a SMF but the mode field diameter, MFD, is a cross-sectional dimension corresponding to 2ω0. However, this
value is hard to measure and the manufacturer-specified value is often nominal or im-precise. For example, the Nufern MFD specification for the 630HP fibre is 4.2 ±0.5 µm measured at 630nm, yielding a variation of more than 30% in the beam divergence calculus. Moreover, Kai et al. showed that although a Gaussian mode can be found to dominate the fibre mode, the exact expansion expression should be considered in the calculation of fibre laser beams propagating through a paraxial optical system[37]. To work around this issue, we instead measured the divergence by imaging the beam exiting from the SMF in a water solution, as depicted in Figure2.3. It yielded an angle of 5.0◦. The angle was fitted into the model using the water refractive index instead of
coreless fibre to retrieve ω0 = 1.30 µm.
The second modelling step is to propagate the beam through the GRIN fibre. To this end, a radially varying refractive index medium, or duct, according to the definitions
2.6 is defined within the ray transfer matrix analysis method n(x) = n0− 12n2x2, γ2 ≡
n2
n0
. (2.6)
The non-square-law behavior of the refractive index gradient was a great concern two decades ago[20], but improvements in production technology eliminated the issue. As demonstrated in Figure2.1 d, index profiles of even low-cost telecommunication fibres such as the Corning Infinicor 300 closely match the theoretical parabolic index vari-ation of the ray transfer matrix theory. The extraction of γ in equvari-ation 2.5 is now straightforward and yields γ = 6.22 /mm, whereas a regular GRIN rod lens gradient constant is usually around 1. A strong dependancy of the GRIN element length over
Intensity (0-255) 100 150 200
Width (pixels)
90 100 110 120 130
Typical Intensity Profile and Gaussian Fit
a) b) Position fr om fibr e edge (µm) 0 5 10 15 20 25
Beam Width 1/e field (µm)
0 20 40 60 80 100 120 140 160 180 200 220 240 260 280 Beam Divergence Angle For Two Different Fibres
3.03992 1 + (x/32.607)2
3.33205 1 + (x/34.9026)2 c)
Figure 2.3: Single-mode fibre output divergence. a) Image of the diverging beam in a water solution. b) Intensity profile along the dotted line in a). c) Plot of the beam diameter against the distance from the fibre edge for two fibres.
the focal distance was therefore expected and observed (see Chapter3.) The equivalent focal length can be calculated using the equation
f = n 1
0γ·sin(γL) (2.7)
and gives f = 100 µm for a 200-µm long piece in air.
The last feature of the probe is a cylindrical lens at the interface between the side of the fibre and the sample medium. This effect has been included in the model but is usually insignificant. Focal spot measurements, performed in water and presented in Chapter 3, exhibit 1–2 µm difference between the focal position in x and y due to this lens. The higher refractive index of tissue further decreases its importance. In brain grey or white matter, the equivalent focal distance calculated with the thin lens approximation is found to be 1 f ≈(n2− n1) 1 R1 − 1 R2 = (1.37 − 1.48)62.51 = −0.00176 (2.8) f ≈ −568 µm. Refractive indices are from coreless fibre measurement and an average of human white and grey matter (Table2.2). R is the fibre radius. From there, the beam is propagated in a homogenous medium of index nt = 1.37 with the same formalism.
Table 2.2: Modelling sample parameters
Parameter Symbol Units λ Human tissue
Refer-ences nm grey Mat-ter White Matter Refractive index nt - 456-1064 1.36 1.38 [58] Scattering Coeffi-cient µs cm−1 640 90±30 400±90 [62,74] Anisotropy g - 640 0.89±0.04 0.84±0.05 [62] Reduced Scatter-ing Coefficient µs’ cm−1 633 27±2 91±5 [70] Absorption Coef-ficient µa cm−1 640 0.2±0.3 0.8±0.3 [62,74]
2.1.2
Optimal Working Distance
The illumination beam simulation is a valuable tool for probe production planning, i.e. to know what lengths to target in order to obtain the desired working distance (WD).
Beam W aist (µm) 0 2 4 6 8 10 Pr obe W orking Distance (µm) −200 0 200 400 600 800 Spacer 1 Length (µm) 100 200 300 400 500 600 GRIN Length (µm) 50 100 150 200 250 300
Figure 2.4: Illumination trends: Working distance and beam waist dependencies on spacer 1 and GRIN lengths. GRIN length was set to 150 µm for variable spacer and spacer length was set to 400 µm for variable GRIN length.
The exact fit between our model and experimental data is discussed in Chapter3, but two trends are observed and highlighted here. The first is the negligible contribution of the first spacer length on focal spot positioning. The top left curve on Figure 2.4
shows an exponential decay of the working distance with increasing spacer length. By choosing a longer spacer corresponding to the tail of the first curve, the effect of length uncertainty caused by the production process (discussed in section2.2.1) is mitigated. The potential resolution is also enhanced. The second finding is that varying the GRIN length leads to the same exponential decay, but has a much greater impact on the focal spot position. The tail cannot be exploited as it causes the focal spot to lie inside the probe. Probes having a GRIN lens out of the 75 − 250 µm interval are to be rejected.
The optimal working distance is a compromise between sampling of untouched tissue and photon propagation in tissue. On the one hand, the working distance must be long enough so that the imaged portion of the sample is never in direct contact with the protective tubing surrounding the probe. On the other hand, the longer the propagation
c) Attenuation Map d) Signal Collection Map b) Collection Map a) Illumination Map
e) Actual Focal Spot Aspect Ratio
Figure 2.5: Modelling intensity maps. The signal collection map (d) is a convolution of the illumination map (a), the collection map (b) and two times the attenuation map (c). a-d are plotted from 0 µm to 180 µm along the propagation axis horizontally and from -4 µm to 4 µm vertically. e) shows the actual aspect ration of the focal spot showed in d). Modelling parameters are 400, 160 and 125 µm for Lspy,Lgr and Lspy respectively.
Tissue parameters are grey matter: µa = 0.2cm−1 and µ0s = 27cm−1.
distance in tissue is, the greater are the effects of scattering and, to a lesser extent, absorption. The straightforward conclusion is that the focal position should be situated right outside the reach of the probe. However, to image completely naive tissue, one should look another 100 µm further[13]. As it is currently, our most compact probe protection has a 300 µm outer diameter, whereas the fibre outer diameter is 125 µm. The working distance must then be of at least 87.5 µm and preferentially over 187.5 µm. The design was nonetheless directed towards a 50 µm working distance to benefit from a higher signal range during preliminary testing.
2.1.3
Illumination and Collection in Tissue
The beam diameter profile is convenient for probe production, but it only takes marginal rays into account. By calculating illumination and collection intensities at multiple points lying on the same transversal plane, a more complete situation is pictured. Assuming rotational symmetry over the beam propagation axis, spatial information can be extracted.
waist dimension ω. ω has been evaluated for every tenth of micron along the propaga-tion axis. Every time, a normalized gaussian illuminapropaga-tion profile was set according to equation 2.9: I = a · exp " −2r 2 ω2 # , a= 2 ω ·√2π. (2.9)
The latter was evaluated 200 times from radial r values ranging from 0 to 20 µm. Figure2.5 b displays the collection of homogeneously and isotropically emitted flu-orescence over the same surface. At each point, a series of rays is launched towards the probe at different angles to define those meeting the collection limits of the confocal system at the SMF interface. The radial collection limit is ω0 = 1.3µm and the angular
limit is derived from Gaussian beam properties θ 'arcsin λ0
nπω0
!
. (2.10)
The algorithm assumes a continuous angle range and transforms this range in a sym-metrical solid angle to estimate the collected fraction of the total fluorescence emitted at this point. Off axis, the real solid angle is not a perfect cone: Its base should rather look like a non-symmetric ellipse. This effect is neglected, resulting in an overvaluation of the solid angle for large radially distanced points.
Finally, attenuation and scattering are taken into account through a third map. Monte Carlo simulations have not been implemented considering that the likelihood of scattered photons to be collected is very low in a confocal system. Instead, the absorp-tion and reduced scattering coefficients are summed up to form the total attenuaabsorp-tion coefficient µt = µa + µ0s. Utilization of the reduced scattering coefficient is legitimate
because the brain is dominated by scattering. Figure 2.5 c shows an attenuation map in human grey matter (µa = 0.2cm−1 and µ0s= 27cm−1).
Fluorescence signal is generated all along the illumination beam propagation axis, and its intensity varies linearly with illumination intensity. The multiplication of the illumination and collection maps represent this situation. To retrieve the effective col-lected efficiency on the x-y surface (figure2.5d), the attenuation map is also multiplied twice: once for illumination propagation to the point and another time for fluorescence propagation to the fibre. The resulting surface is integrated and the normalized collec-tion factor, 4% in this case, is used to quantify probe efficiency. This factor is useful in probe design for inter probe comparison, but does not represent a physical parame-ter as it does not take fluorophore’s quantum yield and fluorescence cross-section into account.
are characterized and are confronted to our model.
2.2
Probe Assembly
Optical fibers are initially juxtaposed by sequential fusion splicing and cleaving on a modified filament fusion splicer. Their end are then cut at 48◦ with a homemade laser
ablation system. After, probes are coated with at least 100nm of aluminum on the angled side to ensure total internal reflexion. Their optical properties are measured and units not meeting optical transmission criteria are discarded. Finally, a permanent FC/APC connector is affixed to the proximal end of the remaining probes for further use on the rotary microscope. Details on each process are provided in this section.
2.2.1
Fusion Splicing
Fusion splicing is a way to permanently join two optical fibres, thus allowing light transmission with minimal scattering and back-reflection at the splicing site. The fibre ends are melted as they are brought together by heat irradiated from an electric arc, a laser, a gas flame or a current-heated filament. The resulting interface is continuous and almost as strong as the virgin fibre. The fusion of optical elements for PIM is carried out at the Centre d’optique, photonique et laser (COPL) facility, using a FFS-2000 Vytran splicer. The arc-fusion splicing system is designed to strip, cleave, splice and recoat special types of fibre. It uses a loop-shaped tungsten filament to provide uniform heating power. Fibres are transferred from station to station while being kept in the same holding block (FHB) at all times. This feature, combined with precise stepper motors and a high resolution CCD camera is key to produce microoptical probes. Normalization routines are executed at the beginning of every splicing session and the FHB motors reset after each probe assembly. The fusion splicing of a single probe is a 15–20 minutes process. Counting all the preparation and handling, 20 probes are typically produced per day. Figure 2.6 a schematizes the result of the fusion splicing process.
Fibre preparation precedes the splicing process. The SMF is at first placed in the left fibre holding block, with roughly four centimeters out on the right side. Its coating is softened by heating on the dedicated stripping/cleaving station and mechanically removed. The fixed Vytran mechanical stripping tool removes slightly less than 2cm of coating once the cleave made. A manual stripping tool is required to remove a
2cm SMF NCF GRIN NCF 125µm 250µm a) After Fusion Splicing SMF NCF GRIN NCF 48° b) After Angle Cleaving 2cm 10,5cm c) After Connectorization
Figure 2.6: Probe assembly schematics after a) fusion splicing, b) angle-cleaving and c) connectorization.
few millimeters more coating in order for the final probe to emerge of the 140– µm inner diameter protective needle. The other fibres do not need extra stripping. Next, the SMF and NCF are cleaned with isopropanol and cleaved using the tension and scribe method. Once the preparation completed, fibres are transferred on the splicing station. On the Vytran FFS-2000, the splicing process is fully automated and highly customizable. Because the construction of microoptical probes imply the uncommon fusion of different fibre types together, custom .xml "Splice files" had to be developed. They correspond to SMF-NCF, NCF-GRIN and GRIN-NCF splices. Parameters such as the filament power, position offset and firing time as well as fibre motion patterns during splicing have been customized to find the perfect recipe for uniform and solid splices. Critical parameters are noted in appendixB.
After a splice, the right FHB is released and the left FHB holds the future probe. The next and most crucial step of the splicing process is to position the fibre for precise cleaving. The spliced interface is centered on the CCD image by a forward and a subsequent backward movement of the left stepper motor to avoid backlash. The left FHB is then backed according to the desired optical element length. The stepper motor can be homed without translating the fibre thanks to a custom secondary holder (modified from Thorlabs, HFV001) installed on a non-moving part of the left FHB. Homing the stepper motor is critical to achieve high-precision cleaving. To further improve length repeatability, the left FHB is pushed towards the cleaving blade with a force of 0.2Nm with the help of a force gauge during the cleaving process. The GRIN element is added following the same routine and since the length of the second NCF is not critical as it will be angle-cleaved, it is left over 200 µm.
Measur ed Element Length (µm) 0 100 200 300 400 500 600 700 800 900 Number of steps 100 150 200 250 300 350 400 450 500 550 600 650 700 750 800 850 Spacer 1 GRIN
Fit over Spacer 1 and GRIN elements
Figure 2.7: Optical Element Lengths and Number of Steps Interdependence Two lengths are critical during the splicing process. The number of steps necessary to move the fibre by the appropriate distance to obtain the desired fibre length has been determined empirically from the assembly of 85 probes. The equation
Number of steps = 1.15 × Desired length(µm) − 153.4µm (2.11) was derived from a linear fit on Figure2.7. The measured amplitude of a step, 1.17 µm, is in good agreement with the correction factor of 1.15 applied to the desired length in microns. The offset is interpreted as the initial distance to pull the fibre back in order to reach the cleaving blade. Length repeatability would greatly benefit from visual feedback during the cleaving process. Adding another CCD camera over the cleaving blade is proposed to this end.
2.2.2
Angle Cleaving
Probe distal ends are angle-cleaved using an optical fibre micro-machining tool custom built at the COPL[27]. The system is composed of a CO2 laser as a heat source, an
acousto-optic modulator to control pulses duty cycle and scanning mirrors. Samples are positioned under the laser beam thanks to a computer-controlled stage and a CCD camera. Silica from the NCF is ablated to form a 48◦±1◦ angled-cleave in a matter of
a) b)
d) c)
Figure 2.8: Angle-Cleaved Probes Under a Scanning Electron Microscope. Images of test probes 7 (a) and 3 (c-d) laser-cleaved surface at different magnifications. Protu-berances are made of dirt, not silica. The prove to be essential for focussing on panels c) and d). Scale bars are located within the images.
The surface makes for a perfect mirror because the glass melted locally during the process. Imperfections on the surface obtained by laser ablation are below the wavelength in size. One can also observe a slight curvature of the cleaved surface on the electron microscopy image (figure2.8). The effect of a slightly "concave mirror" was neglected in the model. Another side-effect of laser ablation is to round the fibre edges. This feature makes the probe less prone to chipping, helps insertion in the needle and may reduce damage to the tissue.
2.2.3
Optical Testing
The optical testing conducted at this stage is preliminary. Its first purpose it to dis-card optically aberrant units. The second goal is to mount the remaining probes on the deposition stage for the upcoming coating step. More precise characterization is described in the next chapter.
A temporary bare fibre adapter is installed on the pre-stripped probe proximal end. 632-nm or 594-nm light is injected in a fibre patchcord which is mated to the probe. It
Side-view Imaging Far-field Imaging Discarded probes Probe 79 Probe 74 Selected probe Probe 86
Figure 2.9: Far-Field and Side-View Probe Imaging. Far-field images have been ac-quired with a HD webcam while the fibre is plunged in water. Probes lay in a diffusing solution for side-view imaging.
is then plunged in a water bath and the projected image of the illumination pattern is qualitatively evaluated, as shown in figure2.9for 3 probes. In theory, the far field image corresponds to the Fourier transform of the beam shape at the focal point. The quality of the focal point is approximated by the projected pattern symmetry and ellipticity. One can easily conclude from far-field images that probe 79 on the right panel is not a good candidate.
The case of probe 74 on the center panel is more litigious. Far-field imaging provides information about the focal spot quality, but not on its position. To address this issue, the illumination beam profile is perceived by its diffusion in a mixture of milk and water with a CCD camera and a 20X air objective. As seen on the lower panel of figure 2.9, probe 74 is rejected since its focal point is situated within the optical fibre. Probe 86 passes both tests and can be mounted with heat-resistant tape on an aluminum plate, along with a dozen others. The illumination beam is used to ensure the cleaved surface faces down.
2.2.4
Coating
The purpose of the metal coating is to ensure total internal reflection on the cleaved face. The aluminum plate supporting the fibre batch is placed in a metal deposition chamber after being cleaned by isopropanol immersion (see AppendixCfor details). At least 100nm of aluminum is deposited. An intermediate chrome layer could be added to facilitate aluminum adhesion to silica although coating durability have not been found to be a major issue. Any other reflective metal in the visible spectrum such as gold could also replace aluminum.
2.2.5
Connector Installation
Having a final probe of 10.5cm – 11.5cm is of paramount importance to fit in PIM. At this stage, the probe is over 15cm long, and the final length is reached during connector installation, which consists of two major steps: assembling the fibre with the connector and polishing the end face. FC/APC terminators were chosen for their low insertion loss and their resistance to vibration. Besides, their angled-polished end face limits back-reflexion. A square snap-on connector is another valuable option.
Probe Experimental
Characterization
Following the ray-tracing excitation model development presented in section 2.1.1, the construction of more than a hundred probes has been initiated. The manufacturing technique evolved over the last two years: The first 20 probes were not angle-cleaved or were mechanically polished, proper metal deposition has only been achieved on the last 40 probes and many of the early prototypes project their focal spot inside the second spacer. To reflect the actual probe capacities, we only considered the most recent probes in this chapter. Their measured properties are listed in table 3.1. A complete table listing the status of every probe that has ever been in production is available on cafeine.crulrg.ulaval.ca[17]. This chapter describes the working distance and resolution measurement methodology. Gathered data is also compared with modelling.
3.1
Working Distance in Tissue
The focal point position can be estimated with the siview imaging technique de-scribed in section 2.2.3, but a more precise method is to rotate the probe and image the incoming beam directly on the CCD chip (the en-face technique). In this case, the 20X air objective is replaced by a 40X and 0.8NA water objective, enhancing resolution. The probe is laid on a translation stage (Sutter Instrument, MPC-200) and bathed in distilled water. The working distance is obtained by measuring the stage travel be-tween the fibre edge and the beam thinnest point, both detected by the experimenter’s eye. The measurement uncertainty is about 1µm due to the experimenter’s ability to
Table 3.1: Measured Probe Properties
Prob
e
No.
Physical dimensions Focal spot characteristics Element length Tip
angle Dimensions Position SP1 GRIN SP2* X Y WD Angle µm ◦ µm ◦ 72 514 1 343 47,0 – – – – 73 378 246 325 47.2 3.47 3.50 -22 6.8 74 545 224 217 47.6 2.54 3.14 -6 6.9 75 513 200 266 47.6 3.89 5.68 31 8.1 76 529 181 276 48.3 4.52 4.35 33 7,0 77 523 180 238 48,0 3.27 4.07 46 8,0 78 543 184 273 47.6 2.99 3.76 41 8.2 79 520 186 285 48,0 – – – – 80 540 195 225 47.8 3.17 3.46 30 7.6 81 522 174 280 48.2 3.58 3.85 69 7.6 82 517 174 281 48.1 5.28 5.66 60 9,0 83 513 202 280 47.8 3.09 3.69 21 7.1 84 526 174 275 48,0 3.76 3.71 67 7.1 85 529 187 278 48,0 3.96 4.11 30 6.2 86 505 161 269 47.7 4.39 5.32 78 8.4 87 526 178 273 47.8 3.23 3.97 51 7.6 88 523 198 278 48.4 3.06 3.56 27 8,0 89 526 161 282 48,0 4.96 6.42 86 8.5 90 528 196 282 48,0 – – – – 91 526 59 277 48,0 – – – 7.4 92 520 99 268 48,0 – – – – 93 521 135 283 47.8 – – – – 94 510 187 278 47.8 – – – – 95 527 222 273 47.8 3.19 3.23 -2 7,0 96 529 295 235 48,0 3.03 3.60 -89 7.3 97 480 160 230 48,0 4.78 5.56 90 7.4 98 474 118 219 47.2 – – – 7.3 99 471 76 242 47.4 15.94 15.53 321 7.1 100 487 225 243 47,0 3.12 3.19 -6 5.8 101 536 172 236 47.4 4.21 4.49 46 7.6
SP: Spacer, GRIN: GRaded-INdex and WD: Working Distance. SP2 refers to the splicing length, later cleaved at roughly 125µm. Dashes denotes inability to perform measurement.
73 74 75 76 7778 80 81 82 83 84 85 86 87 88 89 95 96 97 99 100 101 73 74 757776 8078 81 82 83 84 85 86 87 88 89 95 96 97 99 100 101 Correlation : −0.9892 Fit function: 276.6 - 1.248*x Measur ed W orking Distance (µm) −150 −100 −50 0 50 100 150 200 250 300 350
Spacer 1 Element Length (µm)
360 380 400 420 440 460 480 500 520 540 560
GRIN Element Length(µm)
50 100 150 200 250 300
Figure 3.1: Relation Among Working Distance and Element Lengths
accurately detect the fibre edge. The angle in the beam produces a systematic under evaluation of the working distance of less than 0.01%. Also, the focal spot is elliptical; the thinnest point vertically and horizontally is not on the same plane along the prop-agation axis. The typical difference between the X and Y working distances is 1–2 µm and an average is presented on table 3.1.
We produced this probe subset with the intention to cover a wide range of working distances. Most probes lay between the 0–100 µm interval, but two outliers are worth mentioning: Probe 99, with a short fibre lens has a 321-µm WD. On the other side, probe 96 displays a 89-µm working distance. The latter measurement, along with all the negative WD, may be somewhat off because of the refractive index shift at the probe interface.
Our model in section2.1 predicted a high correlation between the working distance and the fibred lens length, but not with the first spacer length. Figure3.1confirms this postulate: The GRIN length is the most important factor determining where the focal spot will lie. The linear fit on the right panel is only valid in the 150–300 µm GRIN length range because the curve should deviate towards infinity as the lens disappears (probe 99, with a GRIN length of only 76µm, demonstrates this assertion) and increase again for values over a quarter-pitch length. The quarter-pitch length is defined by the waist being right on the lens edge. For longer lenses, the waist lies inside the GRIN fibre and the output beam diverges.
73 74 75 76 777880 81 82 83 84 85 86 87 88 89 95 96 97 99 100 101 73 97 867583 8482817788878995768596 807874 99 100 101 Fit finction: 14465x−1.582 Correlation: −0.9454 Measur
ed X Focal spot size (µm)
0 5 10 15 20 25 30 35 40 45 SP1 Element Length(µm) 360 380 400 420 440 460 480 500 520 540 560
GRIN Element Length(µm)
50 100 150 200 250 300 74 75 76 77 8078 81 82 83 8485 86 87 88 89 9596 97 100 101 73 74 75 76 7778 80 81 82 83 84 85 86 87 88 89 95 96 97 100 101
Figure 3.2: Relation Between Lateral Resolution and Element Lengths
3.2
Resolution
Probe lateral resolution measurements are performed using the same experimental setup as the one used for working distance measurements. An image is acquired at the X (perpendicular to the fibre axis) and Y thinnest dimension and a gaussian curve is fitted on the profile. The 1/e2 intensity point is used to quantify the focal spot size.
The Y resolution is systematically worst than the X resolution. The cause of this difference is astigmatism due to angle cleave.
The link between resolution and working distance is intuitive because it depends directly on the probe numerical aperture. It is therefore normal to observe a relationship between GRIN length and resolution in Figure 3.2. Similarly, the fitted curve should diverge to infinity as the GRIN element shrinks, which is observed. A minimal spot size at the GRIN quarter-pitch length and a step to infinity afterwards is expected as the focal spot enters the lens. This effect takes place beyond the 300 µm GRIN length and is irrelevant here since such a probe always has a negative working distance. As expected, the first spacer length does not affect the resolution significatively.
The axial resolution has not been systematically measured on every probe. It has rather been estimated from beam diffusion images to be around 30 microns. A more accurate measurement is indeed possible with the en-face configuration, but the imaging system should be automated to minimize the time to identify the right plane. A simple but costly solution is to replace the actual CCD camera by a beam profiler.
73 74 7580 787677 81 82 83 84 85 86 87 88 89 95 96 97 100 101 Correlation : 0.9907 Fit function: 32.62 + 0.9364*x Measur ed W orking Distance (µm) !100 !50 0 50 100
Modelled Working Distance (µm)
!100 !50 0 50 73 74 75 76 77 78 80 81 82 83 84 85 86 87 88 89 95 96 97 100 101 Correlation : 0.8018 Fit function: !0.02651 + 3.534*x Measur ed Lateral Resolution (µm) 2.5 3.0 3.5 4.0 4.5 5.0 5.5
Modelled Lateral Resolution (µm)
0.7 0.8 0.9 1.0 1.1 1.2 1.3 1.4
a) b)
Figure 3.3: Modelled Against Measured Working Distance and Resolution Such a low axial resolution is generally undesirable in optical microscopy, However, since our probes do not have the ability to move the focal point axially, the sampled volume increases with the Rayleigh range. In the context of cell body counting for sta-tistical analysis over time, having a poor axial resolution becomes an advantage. Images acquired in a uniformly distributed fluorescent beads solution (section5.2) demonstrate the latter statement feasibility.
3.3
Model Validation
The measured working distances are plotted against modelled values for the same probe geometrical properties and the same medium refractive index in figure 3.3 a. The working distance of actual probes is only predicted accurately if an offset of 33 µm is juxtaposed. This offset is likely due to the GRIN refractive index profile being perturbed because of glass melting on spliced ends, decreasing the effective lens focussing power. This curve fitting is not included in the model and needs to be confirmed by future probe production. With its inclusion in the model, we can expect a few microns accuracy in the working distance prediction.
The values of modelled and measured resolution are less correlated than working distance ones. Variability in the splicing process, such as the cleaving angle, contribute to resolution degradation although they have no effect on the final WD.
It is also noted from the graph 3.3 b that the model greatly overestimates the probes resolution. Considering that the minimal beam diameter provided by the model assumes a purely gaussian beam and no aberrations, the modelled values are rejected.
To put this in context, the numerical aperture of a 100-µm working distance probe is roughly 0.17, which yields a theoretical diffraction-limited resolution of 2 µm according to equation1.2. The model could be improved or other software such as ZEMAX (which supports non-gaussian beams) could help to upgrade the modelled beam propagation, but to obtain a resolution model was not part of our goal in this project.
