Biomaterial microlasers implantable in the cornea, skin, and blood

Abstract

Stand-alone laser particles that are implantable into biological tissues have potential to enable novel optical imaging, diagnosis and therapy. Here we demonstrate several types of biocompatible microlasers and their lasing action within biological systems. Dye-doped polystyrene beads were embedded in the cornea and optically pumped to generate narrowband emission. We fabricated microbeads with poly(lactic-co-glycolic acid) and poly(lactic acid)—substances approved for medical use—and demonstrate lasing from within tissues and whole blood. Furthermore, we demonstrate biocompatible cholesterol-derivative microdroplet lasers via self-assembly to an onion-like radially-resonant photonic crystal structure.

1. INTRODUCTION

Recently there has been an increasing interest in bio-lasers [1,2]. Lasers have several advantages over fluorescence, including narrow emission linewidth, high coherence, large intensity and highly nonlinear output. These properties have been harnessed for ultrasensitive sensing [3], spectral multiplexing [4] and sub-diffraction microscopy [5]. Examples of biological lasers include hybrid lasers containing cells [6,7] and tissues [8] inside Fabry-Pérot cavities, microlasers inside cells [4,9] and random lasers [10] in tissues [11]. Lasers made entirely from biomaterials were also demonstrated. For example, a distributed feedback laser (DFB) was made from riboflavin doped gelatin [12] and silk [13], and whispering-gallery-mode (WGM) lasing has been achieved in water droplets [14,15] and protein microspheres [16].

We have investigated the feasibility of implanting stand-alone bio-lasers into tissues. With exception of random lasers [11], studies on implantable, stand-alone lasers have been lacking [17]. Here we demonstrate the laser action of polymeric microlasers embedded in the cornea, skin and whole blood. Besides chasing a long-standing curiosity of super humans, such as living lasers, capable of emitting lasers from their eyes, the underlying motivation of this study is for potential applications to sensing and diagnosis. For example, implanted lasers in the biological tissues may provide physiological information [18], such as glucose and temperature, in real time.

An important step towards implantable lasers is to achieve biocompatibility; that is, the lasers should not cause immune reactions and foreign-body responses beyond acceptable levels. In some cases, it may be desirable that the lasers are biodegradable. Most of the current bio-integrated lasers contain some non-biodegradable or bio-incompatible materials. In this paper, we employed transparent polymers poly(lactic-co-glycolic acid) (PLGA) and poly(lactic acid) (PLA). Both materials are already approved for medical use and routinely used in clinics for medical implants, sutures and drug delivery, and recently have been also used to make implantable optical waveguides [19]. Microparticles for drug delivery are made from these polymers [20,21], which already have spherical shape and size suitable for sustaining WGMs. We have also explored photonic crystal lasers made from biocompatible, self-assembling liquid crystal chiral molecules. Unlike synthetic liquid crystals extensively used for photonic applications [22,23], cholesterol derivatives [24], specifically cholesteryl esters, are biocompatible [25,26] and found in the human body. Here we demonstrate that droplets of cholesterol derivatives emit laser light and can be used as temperature sensors.

2. RESULTS

A. Polystyrene bead lasers in the cornea

To demonstrate WGM lasing in tissue, we first used bovine cornea, which is transparent. The refractive index of the cornea is ~1.37–1.38. Commercially available green fluorescent polystyrene beads (Thermo Scientific, Fluoro-Max, 8 μm mean diameter, 18% coefficient of variation) were used, which are inert but not biodegradable. A dispersion of these beads in phosphate buffered saline (PBS) with a concentration 2 ×10⁶ beads/ml was used for injection. Bovine eye globes were acquired from a local distributor and the experiments were performed within less than 10 hours postmortem. A hypodermic needle (30 gauge, 0.3 mm outer diameter) was used for the injection of the bead dispersion into the bovine cornea. The needle was inserted at a shallow angle not to penetrate beyond the cornea (Fig. 1a). The needle was slowly retracted from the cornea while at the same time the bead dispersion was pushed out of the needle with a slight pressure. The beads were deposited along the path of the needle (Fig. 1b). For the pumping of the lasers and the collection of light, a 20×0.45 NA objective was used. The microlasers were pumped by an optical parametric oscillator with 5 ns pulse duration, tuned to 475 nm. The laser beam was slightly divergent at the objective entrance pupil, so that the focus at the sample was located slightly further away from the objective focal plane, producing a 20 μm wide excitation area in the focal plane. The collected light was sent through a dichroic mirror to a camera or an imaging spectrometer (300 mm focal length, 0.05 nm resolution). When a single bead was optically pumped, the output showed lasing characteristics, namely a sharp intensity threshold and narrowband (<0.2 nm) laser lines (Fig. 1c).

Figure 1.

Lasing of polystyrene beads in bovine cornea. a) Site of injection of the bead dispersion. b) Fluorescence image of the injected beads. The out of focus beads are on the eye surface. c) Lasing spectrum of a single illuminated bead (inset). Scale bars, 20 mm in (a), 200 μm in (b) and 10 μm in (d).

B. Biodegradable polymer microbeads and lasing in the blood

To make biodegradable lasers, we have produced PLA and PLGA spheres by standard oil in water dispersion procedure. Poly(lactic acid) (PLA, Hycail CML-PLA, MW of 63,000 ± 12,000 Da) or poly(lactic-co-glycolic acid) (PLGA, Sigma, ester-terminated lactide:glycolide 75:25 with a MW of 76,000 – 115,000 Da) were dissolved in dichloromethane (DCM) at a concentration of 4 wt%. Nile red was added to the DCM solution at a concentration of 1 mM. The water phase was prepared by dissolving 1wt% polyvinyl alcohol (Mowiol 4–88, Mw of ~31,000 Da) in water as a surfactant to stabilize the droplet dispersion. The DCM solution was added at 1% to the water phase and vigorously shaken to produce polydispersed droplets. Droplets were ultrasonicated in an ultrasonic bath at the lowest power setting for 10 min. The dispersion was left for at least 15 h so the DCM evaporated from the mixture, leaving solid spheres. The final concentration of Nile red in solid PLA or PLGA beads was 25 mM. The beads were washed several times with water to remove the PVA, with centrifugation at 2,500 g between the washing steps. Beads created by this procedure had sizes ranging from 10 μm to several 100 μm (Fig. 2a and b). Especially larger PLA spheres had some porosity in the center, but smooth nonporous surface. Upon pumping with a green pulsed laser (frequency doubled Nd:YAG laser at 532 nm with 1 ns pulse duration), clear peaks appeared in the emission spectrum from single beads (Fig. 2c and d), corresponding to WGM lasing. The emission intensity versus input energy clearly shows a threshold behavior (Fig. 2e). The minimum size for lasing was approximately 20 μm, limited by the radiation leakage which is highly dependent on refractive index difference between the laser and the exterior as well as the size. PLA and PLLA polymers have a refractive index of 1.47, lower than polystyrene (1.59).

Figure 2.

Whispering gallery mode lasers made from biodegradable polymers. a) PLA beads doped with Nile red. b) PLGA beads doped with Nile red. c) Lasing spectrum of a single 26 μm PLA bead in water. Bright-field image (top inset) and lasing (bottom inset). d) Lasing spectrum of a single 26 μm PLGA bead in water. Bright-field image (top inset) and lasing (bottom inset). e) Output of PLA as the pump energy is increased shows typical threshold behavior. f) A 40-μm diameter PLA bead in blood. The bead is surrounded by red blood cells. g) Lasing of the same PLA bead and h) the emission spectrum. Scale bars, 100 μm in (a) and (b), 10 μm in (c) and (d), 20 μm in (f) and (g).

The operation of these lasers in human blood was tested. Dispersion of PLA beads in PBS was mixed in ratio of 1:1 with freshly collected whole human blood (Fig. 2f). The mixture was introduced in between two glass slides and imaged within 15 min after blood collection. PLA lasers, surrounded by approximately four times smaller red blood cells, showed lasing (Fig. 2g) with narrow lines in the emission spectrum (Fig. 2h). Lasing in blood was as efficient as in water, meaning that the red blood cells and other constituents of blood did not frustrate lasing. Bio-lasers in blood could enable to diagnosis performed directly in blood [27].

C. Polymer microlasers in the skin

To study the lasing properties of polymer mirolasers in biological tissues (Fig. 3a), we implanted beads dispersed in phosphate buffered saline into porcine skin tissues using a tattoo gun (Fig. 3b). Pig skin was purchased from a local supplier and kept in PBS until use. A dense bead suspension containing almost no water was deposited on the skin surface. A tattoo gun with needles at the end was passed few times across the skin where the beads were deposited. The tattoo gun has needles which are moved fast back and forth by an electromagnet to create desired spatial patterns. The tattoo gun was operated at the lowest frequency of the needle movement back and forth. At the same time the skin was repeatedly stretched so that also larger beads could penetrate the holes formed by the needles. The beads remaining on the surface were removed by washing with PBS. Alternatively, hypodermic needles with appropriate inner diameters could be used for injection of the beads. We examined a single bead implanted at a depth of approximately 100 μm below the skin surface. When it was illuminated by the external laser a clear fluorescence emission was observed, but the spherical shape of the laser could not be discerned due to light scattering of the skin (Fig. 3c). In the emission spectrum from the center of the illumination, clear lasing peaks can be distinguished above the broad background fluorescence (Fig. 3d). The background is largely autofluorescence of the skin tissue. From the WGM output spectrum, the bead size was determined to be 49.8 μm.

Figure 3.

Lasers implanted into skin. a) Principle of operation of a laser in skin tissue. b) Implantation of lasers into porcine skin using a standard tattoo gun. c) Light from a laser embedded approximately 100 μm below the skin surface. Because of light scattering the laser itself cannot be clearly distinguished. (d) Output spectrum showing WGM laser peaks at 615–625 nm superimposed on broad fluorescent background. Scale bar, 50 μm in (c).

D. Cholesterol Bragg onion microlasers

We used cholesterol derivatives to form onion-like, spherical photonic crystal structure in a droplet (Fig. 4a). The periodic structure selectively reflects light in accordance with the Bragg law. The cholesteric mixture was prepared from cholesteryl nonanoate, cholesteryl oleyl carbonate and cholesteryl chloride in ratio of 5:4:3. The ratio was optimized so that the long bandedge matched the maximum emission of the fluorescent dye. The reflection was measured by introducing a thin layer of the mixture in between two glass slides and spectrum of white light reflected from the layer was measured with a spectrometer. Pyrromethene 580 (BODIPY) fluorescent dye was introduced into the mixture at a concentration of 2 wt%. All four components were heated up to 90°C so that the cholesterol derivatives reached the isotropic phase (phase transition at 63°C) and the mixture was mixed for 15 min. The mixture was cooled to room temperature and centrifuged at 10,000 g for 5 min to remove any undissolved dye or other particulates. A small quantity of the cholesterol mixture (~1 μl) was introduced into 0.5 ml of glycerol and stirred with a pipette tip to form droplets. Under crossed polarizers the droplets have a cross structure indicating a good liquid crystal orientation throughout the whole droplet volume (Fig. 4b). The molecules are oriented tangentially in each shell and form a helical twist from the center out towards the surface in all directions. Closer inspection reveals a line extending from the center of the droplet to the surface (Fig. 4c). This is a defect structure which is formed due to topological reasons [28,29]. Namely, a sphere cannot be combed without introducing at least one defect, where the orientation of the molecules in undefined.

Figure 4.

Cholesterol lasers. a) Schematic of a dye-doped cholesterol droplet. When optically pumped, the periodic liquid crystal helix supports optical resonance in the radial direction. b) Cholesterol droplets in glycerol between crossed polarizers. c) A single droplet between crossed polarizers with the defect line visible. Arrow indicates the topological defect line. d) Lasing from a single droplet is observed as a bright spot in the center of the droplet. Weak background light and pulsed laser were used to illuminate the droplet. e) Lasing spectra at 25°C and 39°C. f) Positions of the lasing peaks at short and long band-edge as a function of temperature. The shaded area is the physiological relevant temperature range. Scale bars, 50 μm in (b) and 20 μm in (c) and (d).

When a cholesterol droplet was illuminated with an external pulsed laser, laser emission was visible from the center of the droplet (Fig. 4d). In the emitted spectrum one or two lines are visible, corresponding to band-edge lasing on the short and long edge of the photonic bandgap (Fig. 4e). Which line will be lasing depends on the positions of the two edges relative to the gain of the dye. One of the advantages of photonic crystal lasers is that the lasing wavelength is independent of the droplet size, but only on the periodicity. The standard deviation of the lasing wavelength across different droplets is only ~1 nm. This enables more convenient platform for sensing than WGM, where a reference spectrum should be measured for each bead laser, due to unknown diameter, before the sensing is performed. The cholesteric pitch and therefore the periodicity of cholesteric liquid crystals is highly dependent on temperature. The lasing peaks increase almost linearly with the environmental temperature, with a slope of 2.2 and 2.4 nm/K for the long and short band-edge lines, respectively (Fig. 4f). The lasers in water solution were stable on the shelf for a few days after preparation of the droplets, but completely dissolved in water within few weeks, enabling the biodegradation and reabsorption by the organism.

3. CONCLUSIONS

We have demonstrated biocompatible and biodegradable lasers in the forms of solid microbeads and liquid microdroplets. Solid-state microlasers were suitable for operation both in soft solid tissues and in blood. Liquid crystal Bragg onion lasers had a linear temperature dependence in the lasing wavelengths, offering the possibility of temperature sensing, but can be made also temperature insensitive by polymerizing into solid spheres [30,31]. The periodicity of polymerized cholesteric liquid crystals can be made sensitive to a variety of analytes including metal ions [32,33], amino acids [34] and pH [35], for chemical and biomolecular sensing applications. In this work synthetic dyes were employed as gain material, but other biocompatible materials could be used, for example fluorescent proteins [36,37], vitamins [15] and medically approved dyes (fluorescein and indocyanine green). Availability of biocompatible and biodegradable microlasers made from materials approved for medical use or substance already present in the human body may open new opportunities for light-based diagnostics and therapies [38] as well as basic research.