Cancer is one of the leading causes of death worldwide, accounting for nearly 10 million deaths in 2020 [1]. In Canada alone, over 1.5 million individuals have lived or are living with cancer, with over 230,000 new cases and 85,000 deaths in 2022. It is estimated that 43% of people will be diagnosed with cancer in their lifetime [2]. While incidence and death rates have slowly but steadily declined in recent decades, cancer remains a major concern with an aging and growing population.

Chemotherapy is extremely important in the treatment of cancer; in some instances, it is the primary treatment method, particularly with non-solid tumours like leukemia. However, in many instances, chemotherapy is an adjunctive treatment, used with or after surgery or radiation therapy, as patient relapse or metastasis is common [3]. Several factors go into deciding the therapeutic response to a tumour diagnosis. Tumour type is most obvious, as tumour cells have differing sensitivity, responsiveness, or resistance to various therapeutic agents. L-asparaginase, for example, is known to be effective against lymphocytic leukemias, depriving those cells of asparagine, which is essential to their function and survival [4]. Tumour location can pose specific challenges to treatment, such as the difficulty of delivering chemotherapeutic agents to a relatively avascular region, or across the blood-brain-barrier. High doses can systematically force drug to the desired site, but this also risks high systemic toxicity. Smaller tumours tend to be more responsive to treatment than larger tumours; many chemotherapeutic agents work on cells in active growth cycles, and smaller tumours tend to have a larger fraction of cells in this phase [5]. Smaller tumour burden also tends to correlate with healthier patient condition, which often means the patient is able to better tolerate the toxicity of the chemotherapy.

In general, treatment is limited by the negative effects a chemotherapeutic agent has on healthy cells. The ratio of doses between therapeutic effect and toxicity is known as “therapeutic index.” Some adverse effects due to chemotherapy toxicity include gastrointestinal distress, muscular weakness, alopecia, and loss of sensation. However, more severe side effects include cardiac damage, liver damage, and leukopenia/thrombocytopenia, which leaves patients vulnerable to infection or hemorrhaging, respectively [6].

The goal in development of chemotherapy technology is to increase the therapeutic index of both existing and novel treatments and agents. This can be done by increasing the therapeutic effect of an agent, and by decreasing its toxicity. Stronger chemotherapeutic agents have been in development to combat drug-resistant tumours that often lead to treatment failure or patient relapse [3]; they are designed to more aggressively eliminate a larger percentage of tumour cells. Unfortunately, many of these agents do not make it past clinical trials to market, as their increased effectiveness against tumour cells also corresponds with an increased toxicity to healthy cells, leading to more severe systemic toxicity. In addition to toxicity, the development of new chemotherapy drugs is largely limited by bioavailability – the amount of drug in systemic circulation in the body; bioavailability affects the frequency and dosage required for a drug to reach therapeutic concentrations [7]. Hydrophobicity, degradation under physiological conditions or enzymatic processes, and rapid renal clearance are all factors that can negatively impact a drug’s bioavailability and efficacy. Curcumin, derived from the herb, Curcuma longa (commonly known as turmeric in the culinary world), has been identified to have anti-cancer potential due to its ability to interact with a wide range of molecular targets, including multiple signalling pathways related to cancer growth and progression [8–12]. However, its use is limited by hydrophobicity, low bioavailability. And rapid decomposition at neutral or alkaline pH, or exposure to light.

One subset of the field of nanomedicine focuses not on the development of novel drugs and other therapeutic agents, but instead on the development of technologies to improve properties of existing chemotherapeutic agents. Nano and microscale structures and systems greatly expand the potential for drug delivery through several potential factors. Depending on the delivery system, they can assist in drug stability and circulatory half-life, enhance the viability of normally low bioavailable drugs, and/or facilitate targeted delivery and uptake [13]. As a result, these systems can increase therapeutic efficacy, reduce toxicological side effects, and improve patient compliance [14]. Polymeric NPs in particular have been explored as attractive nanoscale drug carriers. Encapsulation of drugs into these nanostructures improves stability and effective solubility, and polymeric NPs have been shown to be easily synthesized through both “bottom-up” methods (e.g., nanoprecipitation, salting out) [15, 16] and “top-down” methods (e.g., solvent emulsion) [17].

Several synthetic polymers including poly (lactic-co-glycolic) acid (PLGA) and poly (ethylene glycol) (PEG) have been approved for use as drug delivery carriers worldwide. PLGA is biocompatible and biodegradable, and its biodegradation and release rate can be controlled by modifying the ratio between the hydrophobic and hydrophilic units (lactic and glycolic acid, respectively) [13, 18]. PLGA nanoparticles encapsulating either hydrophobic or hydrophilic drugs can be easily synthesized using the single or double emulsion method, respectively, and will slowly degrade via hydrolysis to release its load over a period of weeks to months [17].

Additionally, the interest in natural polymers (chitosan, hyaluronic acid, cellulose, etc.) as drug delivery nanomaterial candidates have grown, due to general factors such as biocompatibility, biodegradability, inexpensiveness, and availability [19–21]. Specifically, chitosan has been investigated as a nanocarrier due to the amine groups along its backbone, which make it the only positively charged polysaccharide, and provide reliable points of conjugation for targeting moieties and other structural modifications [22, 23]. A 2014 study showed that chitosan-coated nanoparticles exhibited significant mucoadhesive properties compared to their uncoated counterparts; it is suggested that this property is related to the electrostatic forces between the cationic backbone and negatively charged groups within mucin proteins [23]; another study showed its potential to accumulate in tumour cells [22]. However, chitosan is limited by its low solubility in non-acidic environments, high viscosity in dilute aqueous acidic environments, limited drug release, and low encapsulation efficiency [19, 22, 24]. Instead, chitooligosaccharides (COS), an oligomer of chitosan that can be prepared by depolymerization of chitosan with nitrous acid or hydrochloric acid, has been explored as an alternative that retains chitosan’s desirable physical and chemical properties, while also having better solubility at relatively higher pH [24]. Control over average oligomer length of nitrous acid depolymerization has been shown to be controllable through the stoichiometric ratio between chitosan and nitrous acid [24–26]. This depolymerization also results in a reactive aldehyde group on one end of COS oligomers, which can be functionalized with a variety of groups or moieties.

The properties of synthetic polymers and natural polymers as nanomaterials compliment each other; the former provides tunability of drug release and simple nanostructure synthesis, while the latter provides customization of functional groups and targeting moieties for improved targeting and properties tailored to specific physiological environments. Recent studies have created hybrid nanostructures of synthetic and natural polymers, with synthetic polymers acting as the core structure and natural polymers acting as a coating, indirectly bound by adsorption and intermolecular forces [21, 23, 27]. Some studies have explored PLGA core, chitosan-coated NPs as potential drug nanocarriers [27, 28].

Reliance on physical adsorption for coating of natural polymers may result in decreased stability and sensitivity to changes in pH, ions, temperature, and/or solvents. In contrast, chemical conjugation via forms of linker chemistry can be more resilient, as well as less reliant on physical and electrostatic compatibility of polymer moieties. Amine-reactive crosslinker chemistry can allow for relatively straightforward and facile chemical conjugation because the reaction is regioselective and occurs under benign conditions (aqueous, room temperature, moderate pH) [29]. NHS ester groups, which can be created by modifying carboxyl groups with EDC [N-(3-Dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride] and stabilized with NHS (N-hydroxysuccinimide), are often used as an amine-specific functional group for linking or labelling [30].

While chitosan has been shown to have potential for passive targeting of tumour cells, various ligands can be used in active targeting. Integrins are receptors found in the extracellular matrix (ECM), and are mainly involved in cell-cell and cell-ECM adhesion interactions [31]. Of these integrins, the α_vβ₃ integrin is among the most important in tumour angiogenesis, and is overexpressed in several tumour types including ovarian and breast cancers [32]. The α_vβ₃ integrin serves as a receptors to many ECM molecules that contain the Arg-Gly-Asp (RGD) peptide sequence; as such, ligands based on the RGD peptide, as well as RGD mimetics, have been shown to deliver targeted therapeutic doses of anticancer drugs like Paclitaxel and Doxorubicin to tumour cells [31, 33–35]. The presence of -COOH carboxyl groups in RGD peptide allows for conjugation to COS via NHS-amine reactive chemistry. This allows COS-coated PLGA-core NPs to be functionalized with RGD peptide for active targeting and drug delivery.

The main objective of this study was to develop a modular nanocarrier system utilising facile assembly and conjugation methods that could used with interchangeable core and coating materials, to accelerate successful development of chemotherapy agents. NPs were assembled through the commonly used emulsion method, with PLGA as the core material, and curcumin as the encapsulated model drug. These NPs were coated with COS and RGD in tandem to promote targeting of cancer cells. Figure 1 illustrates the proposed formation of these NPs Physical properties of coated NPs were compared to bare PLGA NPs, and the formulations were compared in vitro via fluorescent cell uptake study and cytotoxicity assay with OVCAR-3 human ovarian tumour cells. This project hypothesizes that PLGA NPs will successfully encapsulate curcumin to protect it from degradation and provide prolonged release, and that NPs coated with COS and RGD will provide better uptake and cytotoxic effects compared to bare PLGA NPs.

Adipic acid dihydrazide, ammonium hydroxide solution (28.0–30% NH₃ basis), ammonium acetate, chitosan (low molecular weight; 50,000–190,000 Da, 75–85% de-acetylated), curcumin [≥94% (curcuminoid content), ≥80%], dimethyl sulfoxide (DMSO), glacial acetic acid, hydrochloric acid (37%) (HCl), MES (4-Morpholineethanesulfonic acid), methanol, methylene chloride (DCM), N-(3-Dimethylaminopropyl)-N′-ethylcarbodiimide hydrochloride (EDC), N-Hydroxysuccinimide (NHS), RGD peptide (Arg-Gly-Asp), sodium cyanoborohydride (NaBH₃CN), sodium nitrite (NaNO₂), trehalose dihydrate were all purchased from Sigma Aldrich. PLGA-COOH [poly (lactic-co-glycolic) acid] (50:50, and 40 K MW) was purchased from Nanosoft Polymers. Cytiva HyClone™ RPMI 1640 Media with Glutamine, Gibco™ Fetal Bovine Serum (certified, heat inactivated), Gibco™ Penicillin-Streptomycin, Phosphate-buffered saline (PBS, 1X, pH 7.4), Pierce Micro BCA Protein Assay kit, Promega CellTiter 96™ AQueous One Solution Cell Proliferation Assay (MTS) were purchased from Thermo Fisher.

COS was prepared from chitosan through nitrous acid depolymerization (shown in Figure 2) based on the protocols by Moussa [24, 26]; low MW chitosan was dissolved in 0.15 M HCl (2 g chitosan per 100 mL solution), then vacuum degassed to remove oxygen from the solution. 500 mM NaNO₂ aqueous solution was prepared (with degassed MilliQ water, immediately before addition to chitosan solution), which was then added for a GlcN:NaNO₂ molar ratio of 10:1 (to achieve COS with approximately 20 average GlcN units), before being stirred moderately overnight so as not to introduce excess oxygen to the mixture. The mixture was filtered to remove undissolved chitosan and other impurities, before precipitating COS by addition of ammonium hydroxide until reaching solution pH 9. To further increase yield, an equal volume of methanol was added to the mixture, which was then centrifuged for 10 min at 10,000 g and 4°C. The COS was washed (x3) by repeated centrifugation in 50/50 water/methanol, before a final wash step with water. The COS was then resuspended in 10 mL water and lyophilized.

Hydrazide functionalized COS (COS-hydrazide) was prepared by dissolving COS in water, with additions of acetic acid. Concentrated MES was then added to a concentration of 0.15 M, before adjusting the solution to pH 4.5 with NaOH or HCl. 10 M equivalents (to COS) of adipic dihydrazide was added to solution and stirred at RT for 24 h, then 10 M equivalents (to COS) of sodium cyanoborohydride was added to solution and stirred at RT for 24 h. As depicted in Figure 3, the COS aldehyde functional group reacts with the hydrazide under acidic conditions to form a hydrazone conjugate, which is further reduced to a secondary amine bond by sodium cyanoborohydride. The COS-hydrazide solution was filtered, precipitated, washed and lyophilized following the same protocol as COS. COS and COS-hydrazide were analysed with ¹H NMR spectroscopy to confirm average size and presence of aldehyde or hydrazide groups, respectively. ¹H NMR spectroscopy was performed on Bruker 500 MHz spectrometer at 300 K. Samples were diluted in deuterated solvent (D2O) to 10 mg/mL; 5 μL HCl (12 M) was added to 1000 μL COS samples to facilitate dissolution in D2O. NMR spectra was analysed using Bruker TopSpin NMR data analysis software.

Void (without encapsulated drug) PLGA NPs were prepared by emulsion method, based loosely on the works of Rafiei and McCall [17, 36]. The initial emulsion mixture consisted of 1:4 v/v organic to aqueous parts. The organic portion was PLGA (acid-terminated) dissolved in DCM to a concentration of 50 mg/mL, while the aqueous portion was 3% w/v PVA. The mixture was briefly vortexed, then emulsified under ice bath via tip sonication. Tip sonication of NP samples were performed with QSonica Q125 Sonicator. Samples were sonicated at 60% amplitude, at 15 s ON, 5 s OFF pulse, for 90 s. The tip was moved around in sample solution during initial sonication pulse to ensure all the organic portion was fully emulsified. The emulsified mixture was then diluted with 1% PVA solution to 300% of its original volume and stirred for 3 h to evaporate the organic solvent. The NP mixture was centrifuged (15,000 g, 10 min, 4°C) and washed with water (x3) to remove excess PVA, before being resuspended in 5 mL DI water (and 5% w/v trehalose as cryoprotectant if necessary) and frozen for lyophilization. A small aliquot was reserved for measuring NP size and zeta potential (ZP). Particle size/distribution and zeta potential of NPs were respectively measured by dynamic light scattering (DLS) and electrophoretic light scattering (ELS) using Malvern Zetasizer Ultra-Red.

PLGA-COS NPs were formed by chemically conjugating COS-hydrazide to the surface of PLGA NPs using amine-reactive NHS chemistry (see Figure 4). COS-hydrazide was dissolved in a small amount of dilute o-phosphoric acid, before being added to MES buffer (100 mM, pH 5) to a COS concentration of 5 mg/mL. PLGA NPs were added resuspended in this solution (2:1 PLGA:COS w/w), along with NHS and EDC (approx. 20%, 30% w/w PLGA NPs, respectively). The mixture was stirred vigorously overnight before washing excess NHS, EDC, and COS through centrifugation. A small aliquot was reserved for measuring NP size and ZP, while the rest was prepared for lyophilization as previously described with PLGA NPs.

RGD peptide was chemically conjugated to COS amine groups through amine-reactive NHS chemistry. 1 mg RGD peptide was added to 10 mL MES buffer (100 mM, pH 6.5) along with 0.5 mg NHS, and 2 mg EDC. PLGA-COS NPs were added and resuspended in the buffer solution through vortexing (and mild sonication if necessary). The NP mixture was stirred overnight, before being centrifuged to separate NP from free RGD. BCA assay was performed on samples of supernatant to determine binding efficiency of RGD to the NPs. NHS concentration across all samples and standards were kept constant for BCA assay analysis, due to the affect and interference of NHS (by reducing the Cu compound in BCA reagent). BCA and other colourimetric assays were recorded and analysed using Varioskan LUX 3020 Multimode Microplate Reader.

The cell uptake study was designed to assess the effectiveness of the COS-RGD-coated NPs at targeting tumour cells versus bare PLGA NPs. RGD peptides have previously been shown to effectively target OVCAR-3 human ovarian carcinoma cells; thus the cell line was used in this study [37]. To better visually compare treatments, Coumarin-6, a hydrophobic fluorescent dye, was used as a replacement encapsulant to curcumin. Coumarin-6 achieves peak excitation around 457 nm.

Briefly, OVCAR-3 cells were cultured and then seeded in a 96-well plate, letting cells incubate for 24 h at 37°C and 5% CO₂. Coumarin-6 in PLGA NPs, PLGA-COS NPs, PLGA-COS-RGD NPs, or free Coumarin-6 (1000 ng/mL) was suspended in cell medium and added to the wells, incubating for another 24 h under the same conditions. The plates were washed with PBS solution, then cells were fixated with formaldehyde solution (4%), and finally stained with DAPI solution (300 nM) for 5 min. The plates were once again washed with PBS and then imaged with fluorescent microscopy. The DAPI stain and Coumarin-6 each showed up under separate channels. ImageJ was used to quantify average fluorescent intensity of Coumarin-6 within cells to determine relative uptake.

As a further demonstration of the delivery effectiveness of PLGA NPs, OVCAR-3 cells were treated with NP-encapsulated and free curcumin for up to 24 h. The different treatments were compared with varying curcumin concentrations (between 2.5 and 40 μg/mL) and treatment times (between 2 and 24 h). To isolate the effects of the delivered curcumin and the NPs themselves, OVCAR-3 cells were also treated for 24 h with void NPs, at the highest and lowest void NP concentrations equivalent to the loaded NP treatments. This void treatment would help determine if the NPs were contributing to cell toxicity directly, in addition to their delivery of curcumin drug.

OVCAR-3 cells were seeded in a 96-well plate at 5000 cells/well and incubated for 24 h at 37°C and 5% CO₂. Treatment medium was produced by resuspending NPs in cell culture medium; free curcumin treatment medium to assist in dispersion of curcumin, which was insoluble in the aqueous medium. After incubation of cells, old medium was replaced with treatment medium, and the cells were incubated under the same conditions as prior (for 24 h in all non-time-dependent studies). After treatment incubation, cells were washed several times with PBS to remove residual curcumin; MTS assay was then performed with Promega CellTiter 96™ AQueous One Solution Cell Proliferation Assay, following the recommended protocol. Colourimetric results from the MTS assay were quantified with Varioskan microplate reader.

All experiments were conducted in triplicate. Data was statistically analysed by unpaired t-test and one-way ANOVA where applicable. In cases where ANOVA found statistical significance, Tukey’s test was used to determine which groups contained statistical differences. The value of α = 0.05 was used to determine whether results were statistically significant.

Depolymerization was successfully confirmed with ¹H NMR spectroscopy. The presence of the amf end unit is confirmed by the small peak at 5.01 ppm that corresponds to the aldehyde proton (specifically the hydrated form of the aldehyde), as well as peaks at 4.05, 4.15, and 4.36 ppm ¹H spectra are corroborated by NMR data of COS from the Moussa study [24, 26]. By comparing the H-2 GlcN to H-1, 3, 4 or 5 amf proton signals in Figure 5, we can see that the average length of each COS molecule is around 20 GlcN units to every amf end unit. Control of the average length of COS was shown to be relatively replicable; average of triplicate samples of COS using a 10:1 GlcN/NaNO₂ molar ratio showed an average observed length of 21.8 ± 1.8 units. Stoichiometric control was further confirmed by increasing the amount of NaNO₂ and observing a change in average oligomer length; average of triplicate samples of COS using a 10:1 GlcN/NaNO₂ molar ratio showed an average observed length of 10.6 ± 2.4 units.

The successful attachment of adipic hydrazide to COS through the amf aldehyde group is confirmed by the absence of the hydrated aldehyde peak (5.01 ppm) as the reaction site, as well as the presence of peaks at 1.57 and 2.31 ppm that correspond to the aliphatic carbon chains between the hydrazide groups, as seen in Figure 6.

PLGA NPs were successfully synthesized by emulsion method; their characteristics were replicated by maintaining consistent parameters including polymer concentration, organic to aqueous ratio, PVA content, and sonication intensity. Similarly prepared PLGA-COS NPs were also prepared, and both were characterized by size, PDI and ZP (zeta potential) via Zetasizer (results shown in Table 1). Both PLGA and PLGA-COS NPs have a similar size (less than 300 nm average), and a relatively narrow PDI (less than 0.05). The COS coating did not appear to affect the size or PDI of the NPs significantly. The negative ZP of PLGA NPs is mainly due to the carboxyl end groups of PLGA molecules at the surface of the NPs. The positive ZP of PLGA-COS NPs suggests the successful coating of COS to the PLGA NP surface, due to the presence of positively charged amine groups in COS. However, PLGA COS NPs could not be directly characterized via ¹H NMR, as the NPs encountered solubility issues -- likely due to the fact that the PLGA portion is highly soluble in organic solvents while the COS portion is highly insoluble in organic solvents, and requires relatively acidic conditions to facilitate solubility even in aqueous media.

When comparing PLGA NPs with and without encapsulated curcumin (via t-test), there was statistically significant difference in NP size and PDI. Comparisons between samples of PLGA COS NPs with and without encapsulated curcumin did not conclude a statistically significant difference. It is however, important to note that curcumin exhibits some fluorescent properties, which caused interference with both the DLS and ELS techniques employed by the Zetasizer to obtain size and ZP data. Fluorescence filters were able to mitigate interference when determining size, but attempts to determine ZP were highly inconsistent. It is highly unlikely that surface charge is affected by encapsulation of drug.

NP and COS coating stability was further investigated by Zetasizer measurements of size, PDI and ZP over a 48 h period, where NP samples (unloaded) were left in aqueous buffer (PBS 1X, pH 7.2, RT) and measured in the Zetasizer at various time points (shown in Table 2). Statistical analysis (ANOVA, α = 0.05) did not show statistically significant changes to either size or ZP in either group over the full 48 h, suggesting that the NPs themselves are relatively stable over the short term, and that the positively charged COS coating remains adhered to the surface, even when dispersed in an aqueous medium.

RGD coating onto PLGA-COS NPs were confirmed by BCA assay of both the supernatant solution, and the NPs directly. The standard curves (Figure 7), which compared the curves of RGD peptide, NHS, and RGD peptide with a controlled concentration of NHS, confirm that NHS interference with BCA assay is consistent and predictable, and that maintaining the same concentration across samples/standards minimizes potential variance due to its presence. The supernatant samples, which contained NHS from the binding reaction, were analysed against the RGD standards containing 12 μg/mL NHS, and diluted appropriately to match this concentration. The NP samples, which were centrifuged and washed of residual reaction reagents, were analysed against RGD standards without NHS. The assays showed a 12.9% ± 0.4% binding efficiency.

Cell uptake of PLGA NP-encapsulated hydrophobic fluorescent dye Coumarin-6 was monitored at various time points up to 24 h. This uptake was compared to a control group with free Coumarin-6. As shown in Figure 9, free Coumarin-6 (Group A) was unable to enter the cell even after 24 h of incubation. Meanwhile, Groups B-D (PLGA NP, PLGA-COS NP, PLGA-COS-RGD NP samples respectively) all showed faint fluorescent signal, and therefore successful uptake, even after 0.5 h of incubation.

The fluorescent images of the OVCAR-3 cells were further analysed with ImageJ software; average fluorescent intensity of an image was determined by sampling intensity at 5 different cells in each image. Results are displayed in Figure 10. Statistical analysis (t-test) determined fluorescent intensity was not significantly different from 8 h to 24 h in Groups B, C or D, suggesting near maximum uptake had been reached by 8 h. Similar analysis between Groups B, C, D showed little statistically significant difference between the overall uptake after 24 h. This seems to suggest that the coated NPs do not provide significant improved delivery over bare PLGA NPs.

OVCAR-3 cells were treated with curcumin-loaded NPs at varying concentrations from 2.5 to 40 μg/mL (shown in Figure 11). The 3 treatment groups of PLGA NP, PLGA-COS NP, PLGA-COS-RGD NP were compared to each other and a positive control group of free, unencapsulated curcumin. Statistical analysis (ANOVA) of MTS assay indicate that change in cytotoxicity with free (unencapsulated) curcumin is not statistically significant between low and high concentrations of curcumin. Comparison of NP-encapsulated curcumin treated samples via ANOVA and Tukey’s test indicate significant difference in cytotoxicity with increasing curcumin concentration; in the 3 NP treatments, more than 5 μg/mL curcumin is required for a statistical cytotoxic effect, while cytotoxicity does not meaningfully increase beyond 20 μg/mL curcumin. IC50 (half-max inhibitory concentration) for all NP treatment groups appears to fall between 10 and 20 μg/mL; ANOVA between the three NP treatment groups did not show significant difference in their cytotoxicity. Specific statistical analysis of varying concentrations for each treatment group is shown in Figure 12.

OVCAR-3 cells were then treated at various with curcumin-loaded NPs for up to 24 h, at a concentration of 40 μg/mL based on the results from Figure 11. As shown in Figure 13, significant cytotoxicity began to occur after 8 h. At this concentration, cell viability appeared to reach 50% around the 16 h mark. Again, free curcumin does not appear to have statistically significant effect on the cells without the assistance of the NP delivery.

To control for the effects of only curcumin and to rule out cytotoxic effects from the NP itself, OVCAR-3 cells were treated with void NPs over 24 h, shown in Figure 14. Each treatment group had a low and high treatment concentration, corresponding respectively to the lowest and highest concentration of curcumin-loaded NPs used in Figure 11. Statistical t-test analysis of the treatment groups did not find any statistically significant difference between high and low concentration samples. The void NPs themselves do not appear to contribute to the cytotoxicity of the curcumin treatment.

Together, these results, along with the fluorescent imaging of Coumarin-6 uptake indicates that, on its own, free curcumin is unable to affect cell viability, most likely due to its insolubility in aqueous medium. The NPs themselves do not appear to have cytotoxic effects, but are shown to be successfully uptaken by cells; when curcumin is encapsulated in these NPs, the curcumin is successfully uptaken by cells, facilitated by the NPs. Once in the cells, the curcumin exhibits cytotoxicity to the cancer cells.

Though cell studies overall did not show statistically significant differences between treatment groups of PLGA, PLGA-COS and PLGA-COS-RGD NPs, it is possible that an exploration into in vitro studies comparing healthy mammalian cells to tumour cells like OVCAR-3 could yield interesting results. The coating process requires further optimization, but the coating process may assist to some degree in avoiding uptake by healthy cells in addition to targeting tumour cells via their receptors.