(188y) Theranostic Optical Fibers for Tumor Treatment and Sensing

Optical fibers have been successfully made from different materials including fused silica, and polymers, and become a versatile platform for optogenetics and neuroscience. Although substantial progress has been made to engineer flexible and multifunctional optical fibers for optogenetics, the use of such fiber for spatiotemporal imaging and drug delivery to cancers has been rarely reported. On the other hand, while drug delivery and nanomedicine have progressed tremendously, the focus of the field has been primarily creating systemic drug delivery particles for tumor imaging or treatment; whereas some tumors remain difficult for nanoparticles to accumulate via systemic circulation, compared with local drug delivery method.

We propose to establish a flexible polymer/drug integrated optical fiber for both diagnosis and treatment of cancers, with minimum mechanical invasiveness. Based on our preliminary data, our fibers can incorporate various dyes and therapeutic agents, regardless of their hydrophilicity. The fiber can release agents over hours, which can potentially enhance the intratumoral drug accumulation with minimal systemic toxicity. Our approach can potentially allow for repetitive on-demand drug release from fibers, overcoming the dosing obstacle.

In addition, to achieve image-guided surgery via optical fibers, we propose to apply photodynamic therapy (PDT) for the treatment. PDT is a local therapy that uses light to activate a photoactivatable molecule called photosensitizer to produce reactive oxygen species, which causes direct cytotoxicity to tumor cells. The photosensitizer-loaded optical fibers could be excellent tools for both treatment and follow-up the tumor response to PDT.

We previously developed a self-assembly nanostructured polylactide (PLA) / poloxamer membrane via a simple solvent evaporation approach. We applied similar approach to coat the silica fibers alternatively with hydrophobic PLA/dye solution and amphiphilic poloxamer-407 solution (Fig. 1a). This method allows the loading of hydrophobic molecules into the coating. For example, after 5 cycle coating, rhodamine B, a fluorescent dye can be successfully loaded, observed by eyes and fluorescence imaging. Once the fiber contact with PBS, the amphiphilic poloxamer induces assembly to form porous surface observed in SEM (Fig. 1b-c), which enables the hydrophobic drug released from the coating. On the other hand, we can loaded hydrophilic molecules into fibers. Our hollow polymer fibers that are manufactured in our lab can draw aqueous solutions into the fiber via capillary force. To avoid rapid leaking of such solution, we used thermosensitive poloxamer aqueous solution that form gel at 37 ^oC to hold the hydrophilic molecule (e.g., fluorescein) in the fiber before implantation (Fig. 1d).

Notably, we can achieve ~ 10 µg/cm drug loading for both hydrophobic and hydrophilic small molecules. We performed in vitro release studies for both types of fibers. Rhodamine had rapid kinetics within 6 hours than verteporfin (over days), a photosensitizer, due to its better water solubility compared to verteporfin. The hydrophilic fluorescein had similar release profile over days in either PC or PMMA fiber. Furthermore, we conducted transwell assays, which eliminate the mechanical effect, to investigate whether the released dye can be uptaken by cells or have therapeutic effect. The release rhodamine from fibers was imaged in 4T1 breast tumor cells 6 hours after incubation (Fig. 1e). Additionally, we incubated 4T1 cells with transwells containing verteporfin-loaded fibers for 4 hours, and irradiated cells with light (650 nm, 20s). After 12h and 24 hours, significant reduction of cell viability were observed in the treated cells, confirming that adequate verteporfin released from fiber was uptaken by cells and induce phototoxicity (Fig. 1f).

To simulate the drug release process through fibers, we used 4% agarose gels as the tissue phantom for fluorescent imaging study. Fibers were dipped into the rhodamine solution and then inserted into the agarose gel blocks immersed in phosphate-buffered saline (PBS) at 37 °C. Fluorescent images taken by the IVIS imaging system confirmed that the dye has diffused from fibers into the phantom in 2 days (Fig. 1g). Ex vivo intratumoral diffusion experiments were performed to verify the in vitro gel diffusion study. Fibers loaded with the aqueous rhodamine solution were inserted in murine 4T1 mammary tumors (excised from balb/c mice) placed in cell culture media at 37 °C for 3 days. The images recorded by IVIS showed that the rhodamine released from the fibers and diffuse throughout the tumors. Similarly, fibers coated with PLA/poloxamer rhodamine were inserted into 4T1 tumors ex vivo; both IVIS fluorescence images and tumor tissue sections clearly indicated the intratumoral diffusion of rhodamine over 10 days.

Currently the delivery of photosensitizer is through intravenous administration, which leads to low accumulation of photosensitizer in tumors. Our approach provides local administration of photosensitizers, as the fiber could be viewed as a depot to release photosensitizer over days. In summary, our proposed approach allows for loading multiple drugs, either in the coating or inside the fiber pores. In addition, the use light can control photosensitizer or drug dose, which prevents the overdose or inadequate dosing. The use of fiber-optics also enables the immediate monitoring treatment outcome via endoscope. Given the ease of engineering and potential programmability, our devices can serve as a new platform for many diseases or conditions conventionally hard to reach, and can synergistically combine multiple therapeutics, including chemotherapy and cell therapy, and imaging modalities for both therapy and diagnosis.