**Figure 5.** Zeta potential of PAMAM G3.5 (solid line), mPEG-G3.5 (dashed line), and CPT/mPEG-G3.5 (dotted line).

Drug loading efficacy is considered to be essential in the design of any drug delivery system, since it relates more or less directly to the therapeutic activity of the system. In our study, the DLE of the proposed system reached 85.37% ± 7.87%, which was exceptionally high compared to previous studies on a similar system of G3.5 but without mPEG conjugation [26]. Therefore, it is conclusive that the mPEG on the surface of the particles can play a significant role in providing more space for drug loading, and also capping the cavities to keep in the drug after loading.

As shown in the drug release profiles of free CPT and loaded CPT (Figure 6), a burst release up to more than 60% of the free drug was observed in the first hour, whereas the loaded system showed a much more prolonged behavior. This could be seen as an improvement in comparison to our prior formulation, which experienced more rapid release [22]. However, this formulation still has room to improve, since the accumulated drug released after a longer period of time should display a gradual rise to be controlled. Yet, a sustainable release of CPT is essentially desirable, as CPT has a very short initial plasma half-life (only up to 2 h) [6], and is rapidly eliminated through urination. Moreover, after administration, the particles together with their cargos need time to reach their target sites, not to mention that in the case of tumor passive targeting, the EPR effect also takes time to accumulate, due to the low leakage rate as mentioned above. During this time, preventing unwanted outflow of drug out of the delivery particles is necessarily important, not only to minimalize side effects to untargeted tissue but also to ensure the amount of drug reaching the target site. Therefore, the design of an antitumor agent delivery system needs to demonstrate an initial non-leakage time.

**Figure 6.** In vitro release profiles over the course of 24 h of free CPT (round) and CPT from mPEG-G3.5 (triangle).

The cytotoxicity of our synthesized system, empty and loaded (mPEG-G3.5 and CPT/mPEG-G3.5) towards normal cells was demonstrated by resazurin assay with normal fibroblasts (L292 line) (Figure 7C-a). As can be seen, up to the concentration of 500 μg/mL for 48 h, the system showed no toxicity to normal cells, which proves the safety of mPEG-G3.5 systems at the tested concentration. Cell viability assay results of free CPT and CPT/mPEG-G3.5 particles on two malignant cell lines, HeLa and A549, were illustrated by live/dead assay (Figures 6B and 7A) and resazurin assay (Figure 7C-b,c). Within the course of 24 h, free CPT established restricted toxicity towards HeLa and A549 cell lines, with a clear concentration-dependent tendency, while CPT/mPEG-G3.5 has not shown significant toxicity. This can be accounted for by two possible explanations: the mechanism of CPT itself and the release profile of CPT/mPEG-G3.5 [18]. CPT exerts toxicity on cells through binding with nuclear DNA forming DNA adducts, including mono- and di-adducts, and inducing apoptosis. However, it is generally accepted that the 1,2-intrastrand di-adduct is responsible for this antitumor effect, due to the fact that it is specifically recognized by High Mobility Group proteins, and less effectively repaired through nucleotide excision [49]. A previous study showed that mono-adducts accumulate in nuclear DNA over 2 days at least, due to a slow diffusion of CPT to nuclear DNA in cells [50], which may interfere with the cytotoxicity rate of both free CPT and encapsulated CPT in our assay, as lower concentrations of drug may establish a lower diffusive gradient, and thus require more time for sufficient accumulation in nuclear DNA. Moreover, as the above release profile implies, the amount of accumulated CPT after 24 h of release also corresponds to the low death rate in the first 24 h. Further in vitro and in vivo research should be carried out in future studies to gain more detail about the therapeutic efficacy of developed formulations.

**Figure 7.** Cell viability after 24 h in percentage and fluorescent signals of HeLa (**A**,**C**-**b**) and A549 (**B**,**C**-**c**) with CPT (square) and CPT/mPEG-G3.5-CPT (round); Cell viability in percentage after 24 h of L929 (**C**-**a**) with mPEG-G3.5 (white dot) and CPT/mPEG-G3.5 (black dot)

#### **3. Materials and Methods**

#### *3.1. Materials*

Carboplatin (CPT, Mw 371.26 Da) was purchased from TCI (Tokyo, Japan). Ethylenediamine (EDA) and toluene were purchased from Merck (Darmstadt, Germany). Methyl acrylate (MA, Mw 86.09 Da), 4-Nitrophenyl chloroformate (NPC, Mw 201.56 Da), poly(ethylene glycol) methyl ether (mPEG, Mw: 5 kDa), tetrahydrofuran (THF), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC), and diethyl ether were supplied by Sigma-Aldrich (St. Louis, MO, USA). Methanol was received from Fisher Scientific. Spectra/Por®® Dialysis Membrane (MWCO 3.5 kDa, 6–8 kDa, and 12–14 kDa) was purchased from Spectrum Laboratories, Inc. (Rancho Dominguez, CA, USA). All reagents and solvents were used as received without further purification.

#### *3.2. Synthesis of Carboxyl-Terminated PAMAM G3.5 Dendrimer*

Half-generation carboxyl-terminated PAMAM G3.5 was synthesized from the EDA core by a divergent approach that employs two consecutive chain-forming reactions, including Michael

addition reaction and amidation reaction, as reported previously by Donald Tomalia with minor modification (Scheme 1) [31]. The Michael addition reaction between primary amine groups of EDA and excess acrylate groups of MA gives the half generation carboxyl-terminated PAMAM denoted by Gn.5, followed by the amidation reaction between methyl propionate groups of carboxyl-terminated PAMAM with excess EDA to generate full generation amine-terminated PAMAM, designated Gn. Shortly, 20 mL of EDA was added to 150 mL of MA dissolved in methanol and kept under constant stirring (3 h at 0 ◦C and then 2 days at room temperature). Solvent and unreacted starting materials were discarded by rotary vacuum evaporator (Strike 300, Lancashire, PR6 0RA, UK) to collect the core precursor G-0.5. After that, a slow addition of G-0.5 (20 g in 10 mL methanol) to 130 mL EDA solution was performed to obtain PAMAM G0.0. The mixture was stirred at room temperature for 4 days, rotated under vacuum using mixed solvent toluene: methanol (9:1 *v*/*v*), and dialyzed by dialysis membrane (MWCO 3.5 kDa) against methanol to eliminate excess EDA and toluene. Finally, methanol was removed from the obtained product by drying under vacuum. These two reactions reiterated continuously to give the next higher generation of PAMAM dendrimer.

**Scheme 1.** Synthetic scheme of amine-terminated poly (ethylene glycol) methyl ether (mPEG) and mPEG-G3.5 dendrimers. The terminal hydroxyl group of mPEG was converted to a primary amine by using 4-Nitrophenyl chloroformate (NPC), followed by the reaction with ethylenediamine (EDA). Amine-terminated mPEG was attached onto G3.5 dendrimer, utilizing 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC) as a carboxyl activating agent.

#### *3.3. Amination of mPEG*

Terminal hydroxyl group of mPEG was converted to primary amine using the following procedure. Firstly, NPC (0.4837 g, 2.4 mmol) was added to molten mPEG (10 g, 2 mmol) at 65 ◦C and stirred for 5 h under nitrogen atmosphere. The mixture was cooled down to 40 ◦C, followed by the addition of 15 mL of THF and then maintaining the constant stirring for 12 h at room temperature. Next, the obtained solution was added dropwise into diethyl ether for the precipitation, filtered, and dried under vacuum to achieve the powdery form of activated mPEG-NPC. Thereafter, mPEG-NPC (7 g, 1.355 mmol) dissolved in distilled water was added dropwise to EDA solution (1.8 mL). The mixture was stirred at room temperature for 24 h. Finally, the resulting solution was dialyzed against distilled water (dialysis membrane MWCO 3.5 kDa) for 3 days and lyophilized to give amine-terminated mPEG.

#### *3.4. Conjugation of mPEG to PAMAM G3.5 (mPEG-G3.5)*

For the conjugation of mPEG to PAMAM G3.5, synthesized PAMAM G3.5 was mixed with EDC 5% and then added into mPEG-NH2 (6.0 g) prepared in distilled water. After 24 h of stirring at room temperature, the resulting solution was put into a dialysis bag (MWCO 12–14 kDa) and dialyzed for 3 days against distilled water to eliminate the impurities. Finally, purified product was lyophilized to obtain powder form for future use.

#### *3.5. Characterizations*

FT-IR/NIR Spectroscopy Frontier (Perkin Elmer, Waltham, MA, USA) and Bruker AC 500 MHz (Bruker Co., Billerica, MA, USA) were used to analyze the chemical structure of synthesized products. To obtain the FT-IR spectra, samples (1–2 mg) were prepared by mixing them with dried KBr (100–200 mg) and then pressing the mixture into a pellet. The measurement acquired in range of 4000–500 cm−1 with a resolution of 4 cm−1. Gas chromatography (GC) (PerkinElmer Clarus 680) analysis of final product was carried out. Synthesized mPEG-G3.5 was dissolved in diethyl ether as solvent for GC run. TEM (JEM-1400, Tokyo, Japan) with an accelerating voltage of 100 kV was used to depict the size and morphology of the products. Sample at a concentration of 1 mg/mL was placed on a carbon-copper grid (300-mesh, Ted Pella Inc., Redding, CA, USA) and air-dried for 10 min before the measurement. Zeta potential and hydrodynamic diameter of the products were determined by Nano Particle Size SZ-100 (Horiba, Kyoto, Japan). Samples were prepared with deionized water at a concentration 1 mg/mL, filtered (pore size 0.45 μm), and sonicated for 5 min. Each sample was measured 3 times.

#### *3.6. Preparation of CPT*/*mPEG-G3.5 and Drug Loading Capacity*

Briefly, 10 mg of CPT was dissolved in methanol, kept under constant stirring for 1 h to completely dissolve the drug, and added slowly to the 10 mL of mPEG-G3.5. The mixture was carried out at room temperature for 24 h, followed by the removal of solvent under vacuum. The resulting mixture was then put into the dialysis bag (MWCO 3.5 kDa) for the dialysis against deionized water in 1 h. The dialyzed solution was freeze-dried to yield the product in powder form while the total volume of solution outside the dialysis bag that contains unloaded CPT was collected to determine the drug entrapment efficiency. The entrapment efficiency (%) of CPT/mPEG-G3.5 was determined by the high performance liquid chromatography (HPLC) method regarding the initial fed CPT and loaded CPT, which was calculated indirectly from the unloaded CPT.

#### *3.7. In Vitro Release Study*

For the in vitro CPT release study, a 1 mL suspension of CPT/mPEG-G3.5 (CPT content, 0.2 mg/mL) in phosphate buffer medium was loaded into a dialysis bag (MWCO 3.5 kDa). The dialysis bag was immersed into a vial containing 14 mL of PBS (0.01 M, pH 7.4), followed by placing the vials in an orbital shaker bath for temperature maintaining (37 ◦C) and continuously shaking (100 rpm). At predetermined time intervals, the release medium (14 mL) was collected, filtered (pore size = 0.22 μm), and replaced with an equivalent volume of fresh media. After lyophilizing the collected media, the amount of released CPT was measured by HPLC (Prominenece LC-20A, Shimadzu, Kyoto, Japan).

#### *3.8. Cell Viability Tests*

Resazurin assay was performed to investigate the cell viability of synthesized products against HeLa cells, MCF-7 cells, and normal mouse fibroblast cells L929. Cells (1.5 × 105 cells/well) were seeded in a 96-well plate containing 100 μL of Dulbecco's Modified Eagle's medium (DMEM, Gibco, Invitrogen) supplemented with 10% fetal bovine serum and 1% penicillin/streptomycin and cultured for 1 day under humidified atmosphere (98%) containing 5% CO2, at 37 ◦C. Then, the media were discarded and replaced with fresh media containing free CPT (10–100 μg/mL) or CPT/mPEG-G3.5 dendrimer containing equivalent CPT concentration for the next 24 h of incubation. In control wells, cells were treated with medium only and assigned to 100% survival. After removing the supernatant, cells were washed twice with PBS, treated with resazurin (10 μL, 0.2 mg/mL), and incubated for 4 h. The fluorescent signal was detected in a micro-plate reader (Varioskan™ LUX, Thermo Scientific,

Waltham, MA, USA) at Ex/Em 560/590 nm. The relative cell viability was calculated by normalizing the fluorescence intensity of samples to that of control group as following equation:

$$\text{Cell viability } \left( \% \right) = \frac{\left( [\text{Abs}]\_{\text{sample}} - [\text{Abs}]\_{\text{blank}} \right)}{\left( [\text{Abs}]\_{\text{control}} - [\text{Abs}]\_{\text{blank}} \right)} \times 100 \tag{1}$$

In addition to the resazurin assay, cell viability assessment of the products against HeLa and MCF-7 cells was visualized by live/dead staining. In short, cells in each well were mixed with 25 μL of fluorescein diacetate (FDA) (10 mM) and ethidium bromide (EB) (7.5 mM). After 3 min of incubation, cells were rinsed several times with PBS. The results were observed using microscope (Eclipse Ti-E Inverted Microscope System, Nikon, Tokyo, Japan) at excitation wavelength of 485 nm for FDA and 530 nm for EB.

#### **4. Conclusions**

Spherical mPEG-G3.5 was successfully prepared with the size within the optimal range for nanocarriers and nearly neutral surface charge. The drug loading efficiency was highly improved, especially the entrapment efficiency, reaching up to more than 80%. Cell viability assay results with three cell lines correspond with the sustained release profile of the prepared particles. Even though mPEG-G3.5 has great cytocompatibility, the lack of anticancer effect of CPT/mPEG-G3.5 compared with CPT on tested malignant cell lines implied that CPT/mPEG-G3.5 is not an appropriate solution as anticancer therapy.

**Author Contributions:** Investigation, M.N.H., N.H.N., N.Q.T., C.K.N. and D.H.N.; writing—original draft preparation, M.T.V. and L.G.B.; writing—review and editing, D.C.N.; supervision, T.T.H.T.

**Funding:** This research received no external funding.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **References**

© 2019 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (http://creativecommons.org/licenses/by/4.0/).

## *Article* **Development of Pectin-Type B Gelatin Polyelectrolyte Complex for Curcumin Delivery in Anticancer Therapy**

**Fu-Ying Shih 1,†, Ih-Jen Su 2,3,4,†, Li-Lun Chu 5, Xiaojie Lin 6, Sheng-Chu Kuo 7,8, Yu-Chi Hou 5 and Yi-Ting Chiang 5,\***

Received: 11 November 2018; Accepted: 14 November 2018; Published: 17 November 2018

**Abstract:** Curcumin has been proven to be a potent agent in colon cancer treatment. However, its hydrophobicity and low oral bioavailability hampered its clinical application. These limitations could be improved through appropriate formulations such as using polyelectrolyte complexes (PECs). PECs were self-assembled with polycations and polyanions in polar solvents. In this study, a novel pectin-type B gelatin PEC was developed for use in curcumin formulation. At pH 4.0, natural polyanions pectin and polycations type B gelatin spontaneously formed PECs in ethanol/water solution, whereas under mimetic gastrointestinal tract (GI tract) conditions, at pH 2.0 and 8.0, pectin and type B gelatin were electrically neutralized, and the PECs swelled to allow payload release. After being transferred to pH 7.0 condition, as in the colon environment, PECs were internalized into colon carcinomas. Thus, pectin-type B gelatin PECs were successfully prepared, and their constituent ratio and drug-loading process were also optimized. The optimum particle size of the PECs was 264.0 ± 3.1 nm and they could swell as the zeta potential was altered at either pH 2.0 or 8.0. The optimum drug content and loading efficiency were 40% and 53%, respectively. At pH 2.0, curcumin was rapidly released from curcumin-loaded PECs, whereas at pH 8.0, curcumin-loaded PECs showed a sustained-release of curcumin. The bare PECs showed very low toxicity toward human normal cells, whereas curcumin-loaded PECs, after incubation at pH 2.0 for 2 h and at pH 8.0 for 4 h, induced cell cycle arrest and exhibited cytotoxic effect to HCT116 human colon cancer cells, even though these loaded PECs were pretreated with mimetic GI tract conditions. Our pectin-type B gelatin PECs were shown to be a promising oral formulation for curcumin delivery in anticancer therapy.

**Keywords:** polyelectrolyte complex; curcumin; anticancer therapy

#### **1. Introduction**

Polyelectrolyte complexes (PECs) are made from two or more macromolecules carrying opposite electrical charges, which can be ionized and electrically complexed with each other in a polar solution [1]. The polymeric acids (polyanions) or bases (polycations) respond to environmental pH value for their ionization in the pKa interval [2]. Therefore, in the field of pharmaceutics, PECs have been developed for decades to encapsulate drugs in oral-controlled or sustained-release system, in response to the environmental pH of the gastrointestinal tract (GI tract) [3]. Besides, PECs are usually formed in polar solvents, such as water or ethanol, to improve the solubility of hydrophobic drugs in pharmaceutics [4]. Among all developed polymeric materials, natural polymers attract much attention for application in drug delivery system because of their excellent biocompatibility and biodegradability [5]. The commonly used polyanions are polysaccharides, whose side chains are appended with carboxylic acids, such as pectin, alginates, and acacia gum [6]. Proteins, such as zein and gelatin, and modified polysaccharides, such as chitosan, are widely utilized as natural polycationic materials [7–9]. In addition, natural polymers as PECs have been applied in drug delivery systems to improve the water solubility or for controlled release designs in many studies [1]. For example, D. Wurster et al. have prepared natural polymers containing cationic type A gelatin and anionic acacia gum to generate PECs for encapsulating hydrophobic agent vitamin A palmitate. The mixing ratios and fabricating methods have been studied and optimized. The PECs could efficiently improve the water solubility of vitamin A palmitate [10]. B. Stokke et al. have built and optimized PECs with natural polymers containing cationic chitosan (pKa 6.5–6.6) and anionic alginates (pKa 3.4–3.7) at the pKa intervals, and they measured the average diameters upon environmental pH changes from pH 3.0 to 8.0. Moreover, they have found that particle sizes would significantly change at above pH 7.0 for the zeta potential alteration (above the pKa of chitosan), that would be potential in controlled releasing pharmaceutical designs [8].

Curcumin (diferuloylmethane) is a phenolic component extracted from the rhizome of turmeric [11]. As a bioactive compound, it has proven its medical efficacy in several tumors [12–14], including gastric [15] and colorectal cancers [16,17]. However, its low water solubility and bioavailability limit its clinical uses via oral administration [18,19]. Besides, curcumin in a buffer decomposes in a pH-dependent manner, and this degradation accelerates at neutral-basic conditions in preservation, strongly diminishing its pharmacological activity [20]. Studies have suggested that curcumin could be well formulated to improve its solubility and bioavailability via a reliable controlledor sustained-release system [21]. PEC is one of several promising approaches and several attempts to generate PECs for curcumin delivery have been conducted [22,23]. K. Hu et al. have fabricated nano-sized zein-pectin PEC for curcumin delivery [7]. P. Sarika et al. have prepared a cationized gelatin and sodium alginate PEC to improve the water solubility of curcumin, allowing sustained-release of curcumin and enhancing its cytotoxicity in vitro [24]. A. Anitha et al. have prepared curcumin-loaded dextran sulfate-chitosan PECs to increase the water solubility of curcumin, and the curcumin content also exhibited cytotoxicity toward cancer cells [25]. Although the principle of PEC formation and studies have proved that PECs would be an applicable formulation for curcumin, it is still difficult to optimize their manufacturing process, including determination of constituent ratios and pH for aggregation [26]. These manufacturing factors strongly affect the particle size, morphology, loading efficiency, and drug-releasing behaviors of PEC systems [27]. Thus, it is essential to develop a PEC system with an optimized manufacturing process.

In this study, we proposed a novel PEC composed of natural polymers, and its fabrication was optimized based on its performance in mimetic GI environment. We assembled the natural polycation type B gelatin and the polyanion pectin into a PEC system at an appropriate pH in a polar solvent system, as shown in Scheme 1. Several studies have previously reported pectin-gelatin PECs, yet type A gelatin was mostly chosen for use in their systems [26,28]. Type B gelatin, to our best knowledge, was first introduced in PEC for curcumin delivery. Type B gelatin is derived from collagen after lime hydrolyzation, and its isoelectric point (pI) was 4.8–5.5. The cationized type B gelatin would yield

below pH 4.8, such as at pH 4.0. Apart from the electrical properties of type B gelatin, the pI of type A gelatin is approximately 9.0; thus, it is positively charged under this broadline [29]. Importantly, the pKa of polysaccharides pectin is 3.5 [30]. This pectin has been utilized to fabricate PEC because it could be electrically neutralized to force payload release in a stomach environment, which has a pH of 2.0 [31]. Besides, recent studies have pointed out the potential of pectin in colon cancer treatment because its backbone could only be cleaved by colon microorganisms [32]. Herein, PECs composed of pectin and type B gelatin were prepared at pH 4.0, which could reduce the decomposition of curcumin during manufacture or preservation. The curcumin-releasing profiles of the PECs were further investigated in the mimetic GI environmental pH value and transit periods (pH 2.0 for 2 h to simulate conditions in the stomach, and pH 8.0 for 4 h to mimic pH conditions in the small intestine) [33]. The remaining curcumin-loaded PECs after pretreatment with mimetic GI tract conditions would be internalized into the colon carcinoma. The biosafety and cytotoxic effect of curcumin-loaded PECs to human normal cells and colon cancer were examined in this study to evaluate the potentials of PECs as a curcumin formulation for future clinical anticancer treatment.

**Scheme 1.** Preparation of polyelectrolyte complexes (PECs). Pectin and type B gelatin were first prepared into diluted polymer solutions in ethanol/water (1:10, *v*/*v*). At pH 4.0, pectin and type B gelatin were mixed together, spontaneously constituting PECs by their charged side chains.

#### **2. Results**

#### *2.1. Preparation and Characterization of Pectin-Type B Gelatin PECs*

PECs were spontaneously assembled in the vicinity of the pKa or pI intervals of the two constituting polymers. The pI of type B gelatin was 4.8–5.5, indicating that the amine groups appended in type B gelatin would be electrically cationic below pH 4.8, whereas pectin, whose pKa is 3.5, would be anionized at above pH 3.5 owing to the ionization of the carboxylic groups appended on its side chains [30]. To fabricate PECs in this study, we first prepared type B gelatin and pectin into 0.1% and 0.02% *w*/*w* diluted aqueous solutions, respectively, to prevent rapid aggregation and formation of large particles. Next, we adjusted the pH value into 4.0 to allow simultaneous cationization of type B gelatin and anionization of pectin. Opposing electrical charges would spontaneously drive the type B gelatin and pectin into PECs in the polar solvent. Considering the subsequent curcumin-loading process, in this study we used ethanol (10% *v*/*v*) as a polar solvent in our PEC system. Lastly, the cationic type B gelatin, anionic pectin, and ethanol would be adjusted to optimize our PEC system.

To optimize the PEC system, the components (mainly pectin and type B gelatin) were assembled various ratios into the PECs, and particle size and distribution were measured by dynamic laser scattering (DLS). The results are presented in Table 1. Table 1 indicated that particle size significantly decreased along with reduction in type B gelatin content until only 60% of type B gelatin remained the PECs. For example, the particle size of the P1G9 sample (10% mole of pectin and 90% of type B gelatin) was 2745.0 ± 12.8 nm, while that of the P4G6 sample (40% mole of pectin and 60% of type B gelatin) decreased to 237.0 ± 3.3 nm. Because the type B gelatin content decreased by less than 60%, the particle size would not be dramatically altered. However, the polydispersity indexes (PDIs), which represents size distribution, would gradually increase with the continual decrease of type B gelatin content. To further investigate the components' impact on electrical properties and environmental pH, we chose to use P4G6 (40% mole of pectin and 60% mole of type B gelatin) and P6G4 (60% mole of pectin and 40% mole of type B gelatin), owing to their similar particle sizes and PDIs.

**Table 1.** The particle size and distribution of the PECs containing various ratios of constituents (to a final concentration of 1 mL).

#### *2.2. pH Responsiveness of the PECs*

The particle sizes of P4G6 and P6G4 were 264.7 ± 3.1 nm and 243.1 ± 18.8 nm, respective, whereas their PDIs were both approximately 0.23. Their morphologies were observed using transmission electron microscopy (TEM) and scanning electron microscopy (SEM). The TEM images are shown in Figure 1 and they show that the observed particle sizes of P4G6 and P6G4 corresponded to their hydrodynamic diameter, as measured by DLS. However, their morphologies were different. The TEM images indicated that P4G6 particles exhibited spherical structure, whereas P6G4 particles showed irregularly spherical structure, probably as a result from the differences in their components. The morphologies of P4G6 and P6G4 PECs were also observed by a scanning electron microscopy (SEM) after the PECs were freezed dried. The SEM images of P4G6 and P6G4 were respectively present in Figure S1a,b in Supplementary Materials. P4G6 PECs exhibited dense nanosperical structure, while P6G4 PECs was getting loose after being freeze-dried. That further confirmed the components of PECs might affect the PEC structure. Both TEM and SEM images suggested that P4G6 PECs exist the stronger electrical interactions to maintain their structure, even after being freeze-dried. However, both P4G6 and P6G4 PECS could also be observed deformation by the SEM after being freeze-dried, that also suggested our PEC system should be phramaceutically delivered in a solution form. Herein, the following experiments would be perfomed and evaluated as PEC is delivered in the solution form.

By analyzing the TEM images, we identified that the components affected the particles' particular shapes. Responsiveness to environmental pH would be influenced by the PEC components. In this study, we mimicked pH changes in the digesting process of drugs following oral administration. In the stomach, the pH is 2.0 on average and maintained for 2 h; in the small intestine, the pH increases to 8.0, which is retained for4h[33]. Herein, we simulated the pH values and periods of drug or food digestion in the GI tract following oral administration to investigate the pH responsiveness of PECs with respect to particle size, and then optimized the PEC components. The PECs were first incubated at pH 2.0 for 2 h; afterward, the PECs were incubated at pH 8.0 for 4 h. The resulting particle size and distribution are presented in Figure 2. Figure 2a,b indicated the particle size and distribution of P4G6 and P6G4 PECs, respectively, at pH 2.0 and 8.0. The trends in size and PDI alterations were marked. The P4G6 PECs showed a steady increase in particle size and PDI values at either pH 2.0 or 8.0. P6G4 PECs exhibited an abrupt increase in size and PDI at pH 2.0, although these PECs also significantly changed their hydrodynamic diameter at pH 8.0. The different performances of particle alteration could be attributed to the different ratios of pectin and type B gelatin within the PECs, resulting in different electrical properties in various pH conditions.

**Figure 1.** TEM images of the P4G6 (**a**) and P6G4 (**b**) PECs. The PECs were dropped to carbon-coated cooper grids. After the PECs were attached onto the carbon cooper grids, excess PECs were removed. The attached PECs were then stained with 1% sodium phosphotungstate aqueous (PTA) solution. The 1% PTA solution was dropped to the grids and excess PTA solution was removed after staining. The TEM grids were dried in a vacuum oven and stored at room temperature until observation (scale bar: 500 nm).

**Figure 2.** Particle size and distribution of the P4G6 (**a**) and P6G4 (**b**) PECs. The PECs were incubated at 37 ◦C and pH 2.0 for 2 h, and the pH was subsequently adjusted to 8.0 for 4 h to mimic drug transits in the GI tract. At each hour, particle size and PDI were determined by DLS. The results have been statistically compared and analyzed using Student's *t*-test (Microsoft Excel 2000). Differences were recognized as statistically significant when the *p* values were less than 0.05, and significant differences are shown in star marks (\* *p* < 0.05; \*\* *p* < 0.01, and \*\*\* *p* < 0.001).

Figure 3 shows the zeta potentials of P4G6 and P6G4 PECs in different pH environments, which mimicked those in the GI tract. At pH 2.0, P4G6 PECs possessed higher zeta potential than that of P6G4 PECs, whereas P6G4 PECs were more negatively charged than P4G6 PECs at pH 8.0. That may be attributed to the different ratios in the PEC compositions. At pH 2.0, pectin was neutralized owing to the protonation of the appended carboxylic groups of pectin, whereas the type B gelatin still retained its cations [29]. P4G6 PECs, which consisted of more type B gelatin, would be more positively charged. In PECs at pH 8.0, pectin was a polyanion, whereas type B gelatin would be electrically neutralized. P4G6 PECs, which comprised of less pectin, would lead to less-negative charges. P4G6 PECs could respond to the outer pH environment via superficial electrical potentials, which would not result in destruction of PECs in either acid or alkaline condition. Herein, P4G6 would potentiate a drug delivery system with the sustained-release formulation of curcumin for colon cancer treatment.

**Figure 3.** Zeta potentials of the PECs. P4G6 and P6G4 PECs were placed at 37 ◦C and pH 2.0 for 2 h, and the environmental pH was adjusted to 8.0 for 4 h to simulate drug transportation in the GI tract. At the predetermined time points (each hour), zeta potential was determined by electrophoretic light scattering (ELS). The results have been statistically analyzed using Student's *t*-test (Microsoft Excel 2000). Differences were recognized as statistically significant when the *p* values were less than 0.05, and significant differences are shown in star marks (\* *p* < 0.05 and \*\* *p* < 0.01).

#### *2.3. Preparation and Optimization of Curcumin-Loaded PECs*

On the basis of previous results, we have chosen the P4G6 PECs as the optimum formulation. This formulation was subsequently loaded with curcumin and the loading process was also optimized. To load the hydrophobic curcumin, it was first prepared at various concentrations in ethanol, which involved our formulation as the solvent, forming 0.5, 1, 1.5, 2, and 2.5 mg/mL alcoholic solution. The alcoholic solutions were independently mixed with type B gelatin and pectin, following the abovementioned procedures. After encapsulation, the particle size and distribution of the PECs were assessed using DLS, as shown in Figure 4a. In all curcumin-loaded PECs, the particle sizes were less than 270 nm. The PECs made from 1.5 mg/mL of curcumin-loaded alcoholic solution exhibited the smallest particle size, with PDI values of 0.28, which was the highest of all PECs'. Along with the concentration of the loaded curcumin (2 and 2.5 mg/mL), particle size increased; as loading-curcumin concentration decreased (0.5 and 1 mg/mL), particle size and PDI value decreased. By ANOVA, we observed that the curcumin-loaded PECs made from 0.5 mg/mL of curcumin alcoholic solution exhibited the ideal particle size (214.67 ± 3.43 nm) and the PDI value (0.17 ± 0.01), which indicated smaller and more homogeneous particles than those of PECs without curcumin load.

Besides the particle size and PDI of the curcumin-loaded PECs, drug content and loading efficiency should also be considered to optimize curcumin encapsulation. The curcumin-loaded PECs were first freeze-dried. The entrapped curcumin was extracted using methanol from the freeze-dried powder and measured by an enzyme-linked immunosorbent assay (ELISA) reader. Drug content could be

defined as the ratio of entrapped curcumin in all curcumin-loaded PECs, and loading efficiency could be determined by the ratio of loaded curcumin in comparison with the total input curcumin during the loading process. Drug content and loading efficiency are presented in Figure 4b. The highest drug content (40%) and loading efficiency (53%) were both observed in PECs made from 0.5 mg/mL of curcumin alcoholic solution. The PECs made from other loading concentrations of curcumin alcoholic solution contained less than 10% of curcumin, with loading efficiencies of lower than 35%, which resulted in a higher concentration of curcumin and interference of PEC formation [34]. The PEC made from 0.5 mg/mL of curcumin alcoholic concentration could result in not only small and homogeneous particles but also the highest drug content and loading efficiency. Herein, drug-releasing behaviors were further studied using the curcumin-loaded PECs comprising of 40% mole of pectin and 60% mole of type B gelatin, and loaded with 0.5 mg/mL of curcumin alcoholic solution.

**Figure 4.** Drug-loading of curcumin-loaded PECs. Curcumin was prepared at various concentrations in alcohol. The curcumin alcoholic solution was mixed with type B gelatin and pectin, forming PECs at pH 4.0. The particle size and PDIs (**a**) of curcumin-loaded PECs were determined by DLS. To further determine drug contents and loading efficiency (**b**), the curcumin-loaded PECs were subsequently freeze-dried. The powdered curcumin-loaded PECs were reconstituted with methanol. Curcumin concentration in the solutions was measured by ultraviolet-visible light spectrum at 423 nm.

#### *2.4. Drug-Releasing Profiles*

The drug-releasing behaviors of curcumin-loaded PECs in mimetic GI tract conditions (pH 2.0 for 2 h and pH 8.0 for 4 h) were also studied. The drug-releasing profile is shown in Figure 5, where two releasing trends are observed based on the pH values. For the first 2 h at pH 2.0, less than 20% of curcumin was released. As the pH increased to 8.0, the PECs released curcumin sustainably and eventually a total of 60% of curcumin was released. It is worth noting that the drug-releasing curve at pH 2.0 was different from that at pH 8.0. At pH 2.0, the drug-releasing curve was close to first- order drug-releasing kinetics, whereas at pH 8.0, the drug-releasing curve was close to zero-order kinetic drug-releasing profile, releasing curcumin sustainably. Besides, almost 40% of curcumin still remained inside the PECs after 6 h of acid or alkaline treatment. The residual curcumin would be provided to treat colon cancer.

#### *2.5. Confocal Laser Scanning Microscopic (CLSM) Observation*

According to the results of drug releasing profiles (Figure 5), 60% of curcumin would be released out of the PECs in mimetic GI tract conditions. That meant 40% of the curcumin would still remain within the PECs, which has possibility to be transported to the colon lesions and uptaken by the colon cancerous cells. To confirm that the curcumin-loaded PECs could preserve curcumin and be internalized into HCT116 human colon cancer cells even after acidic or alkaline treatments, we used a CLSM for direct fluorescent observation. In order to observe the internalization of PECs to HCT116 cells, the PECs were labeled with cyanine 5.5 monosuccinimidyl ester (Cy 5.5-NHS ester) fluorescence via amide bonds with the amine groups appended on type B gelatin. The organelle endosome/lysosome was marked with Lysotracker DND-Red and cell nuclei were stained with 4',6-diamidino-2-phenylindole (DAPI). Curcumin was detected by its own fluorescence. Curcumin and Cy 5.5-labeled curcumin-loaded PECs were incubated in mimetic GI tract conditions and added to HCT116 cells, which were then incubated for 12 h. The fluorescence of Cy 5.5, lysotracker, curcumin, and DAPI were independently excited and detected with appropriate excitation/emission wavelengths; they are shown in grey, red, green, and blue, respectively, in Figure 6.

**Figure 5.** Drug-releasing profiles. Curcumin-loaded PECs were placed into dialysis bags (molecular weight cutoff: 6–8 k) (M.W.C.O.: 6–8 k), which were immersed to an ethanol/deionized water (1:1, *v*/*v*) solution at 37 ◦C. The pH of the outer medium was first adjusted to 2.0 for 2 h and then adjusted to 8.0 for 4 h to mimic drug transit in the GI tract. At the predetermined time points, curcumin release was determined by UV spectrum at a wavelength of 423 nm.

**Figure 6.** Confocal laser scanning microscope (CLSM) images showing internalization of curcumin (**a**) and curcumin-loaded PECs (**b**). HCT116 human colon cancer cells were incubated with curcumin or Cy5.5-labled curcumin-loaded PECs, which were pretreated with mimetic GI tract conditions. After 12 h of incubation, the cells were stained with Lysotracker DND-Red to label the endosome/lysosomes, and then fixed with paraformaldehyde. The fixed cells were then washed twice with PBS and further stained and mounted with 4 ,6-diamidino-2-phenylindole (DAPI)-containing mounting medium. The fluorescence of the Cy 5.5-labeled curcumin-loaded PECs, curcumin, endosome/lysosomes, and DAPI was observed using a CLSM with appropriate excitation and emission wavelengths.

Figure 6 shows the internalization process of curcumin and curcumin-loaded PECs. Curcumin could enter the cells via diffusion and spread over the organelles within the cells. Herein, in Figure 6, the fluorescence of curcumin could be observed over the cell cytoplasm, even also be observed to overlap with that of endosome/lysosomes. Curcumin-loaded PECs were observed to be internalized to HCT116 cells, and mainly accumulated in the endosome/lysosomes. Curcumin fluorescence was also distributed within the cells, showing the intracellular drug-releasing behaviors of curcumin-loaded PECs. It is worth noting that the fluorescence of the free curcumin was weaker than that of curcumin-loaded PECs, leading probably from the chemical structure degradation of free curcumin. That would be an evidence for the protective effects of our PECs. The PECs would carry active curcumin into the colon cancer cells even through the mimetic GI tract conditions.

#### *2.6. Cell Viability*

In order to verify the therapeutic efficacy of curcumin-loaded PECs against colon cancer cells, HCT116 cells were incubated with curcumin-loaded PECs or free curcumin. The free insoluble curcumin was first prepared in 1% of dimethyl sulfoxide (DMSO), to the same concentration as curcumin concentration in the curcumin-loaded with PECs'. Both curcumin-loaded PECs and free curcumin were pretreated with mimetic GI tract conditions and transit periods, as mentioned previously. After that, curcumin and curcumin-loaded PECs were prepared at concentrations of 45 to 1.41 μM by serial dilution and respectively added to HCT116 cells. The resulting cell viability was determined by (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) (MTT) assay after incubation.

The viability of HCT116 cells following treatment with curcumin and curcumin-loaded PECs, which were pretreated with mimetic GI tract conditions, is shown in Figure 7. The results indicated that both curcumin and curcumin-loaded PECs exhibited concentration-dependent toxicity to HCT116 cells. The half maximal inhibitory concentrations (IC50) of free curcumin and curcumin-loaded PECs were approximately 25 and 5 μM, respectively. The results directly showed that curcumin would be degraded in mimetic GI tract conditions, thereby reducing its toxicity to colon cancer cells. Curcumin-loaded PECs enabled sustained-release of their payloads in mimetic GI tract conditions, and the released payloads exhibited toxicity to cancer cells. Besides, the PECs protected curcumin from degradation due to environmental pH during treatment with mimetic GI tract conditions; hence, more bioactive ingredients were able to arrive at tumorous lesions in the colon. Herein, the curcumin-loaded PECs exhibited higher toxicity toward HCT116 human colon cancer cells than free curcumin.

Besides, the safety of the novel bare PECs was also evaluated using Detroit 551 human normal fibroblast cells, and the MTT assay was employed to determine cell viability. The concentrations of PECs were adjusted on the basis of drug contents until they were equal to those used for curcumin encapsulation. Cell viability data are presented in Figure S2 in the Supplementary Materials. Bare PECs exhibited very low toxicity toward human normal fibroblasts after they were directly cultured with the cells. Even though the bare PECs were incubated after treatment with mimetic GI tract conditions, as described above, the bare PECs exhibited very low cytotoxicity. The results clearly illustrated that our pectin-type B gelatin PECs possessed high safety to human normal cells, and that the bare PECs were digested after acid or alkaline treatments.

**Figure 7.** Cell viability. Curcumin and curcumin-loaded PECs were pretreated with pH 2.0 for 2 h and pH 8.0 for 4 h. Afterward, the pH values were adjusted back to 7.4. Next, HCT116 human colon cancer cells were incubated with the curcumin and curcumin-loaded PECs at 37 ◦C with 5% CO2 supply for 12 h. The curcumin and curcumin-loaded PECs were removed and the MTT assay was utilized to determine cell viability. The results have been statistically compared and analyzed using Student's *t*-test (Microsoft Excel 2000). Differences were considered to be statistically significant when the *p* values were less than 0.05, and significant differences are shown in star marks (\* *p* < 0.05; \*\* *p* < 0.01, and \*\*\* *p* < 0.001).

#### *2.7. Cell Cycle Analysis*

In order to further identify the relevance of the cytotoxicity and protective ability of our PECs from exteriors, HCT116 cells were treated with curcumin or curcumin-loaded PECs, which were both preincubated in mimetic GI tract conditions (pH 2.0 for 2 h and pH 8.0 for 4 h). The cell cycle of HCT116 cells was analyzed using flow cytometry after cell fixation and staining by propidium iodide (PI). The results are shown in Figure 8. Figure 8a shows that curcumin after incubation in mimetic GI tract conditions did not cause cell cycle arrest, whereas Figure 8b indicates that curcumin-loaded PECs caused G2/M phase arrest. The results were consistent with the cytotoxic results, as shown in Figure 7 Curcumin-loaded PECs could maintain curcumin bioactivity from the environment, leading to stronger cytotoxic effects of curcumin on colon cancer cells.

**Figure 8.** Cell cycle profiles of HCT116 cells treated with curcumin and curcumin-loaded PECs after incubation in mimetic GI tract conditions. (**a**) Curcumin and curcumin-loaded PECs (curcumin concentration was 6.8 μM) were pretreated with mimetic GI tract conditions, and HCT116 human colon cancer cells were then treated and the curcumin or curcumin-loaded PECs for 12 h. The cells were then fixed and stained with PI and cell cycle was analyzed using flow cytometry; (**b**) Cell cycle distribution was also analyzed using ModFit LT 3.0. The results have been statistically compared and analyzed using Student's *t*-test (Microsoft Excel 2000). Differences were recognized as statistically significant when the *p* values were less than 0.05, and significant differences are shown in star marks (\* *p* < 0.05 and \*\*\* *p* < 0.001).

#### **3. Discussion**

Curcumin is a potent anti-cancer compound, although its clinical application is limited owing to its hydrophobicity, poor oral bioavailability, and unstable chemical structure. To overcome these challenges, we aimed to design an appropriate curcumin oral formulation for colon cancer treatment. PECs are one of the feasible approaches for achieving this formulation.

PEC is composed of two or more oppositely charged polymers in polar solvents. PEC could not only improve the water solubility of the encapsulated agents but also protect its payload from the environment. Besides, the electrical charge of PEC would change with the environments, thus potentiating oral-controlled or sustained-release drug delivery system using suitable constituents of PEC to enhance oral bioavailability of drugs. In this study, anionic pectin and cationic type B gelatin used to form PEC for curcumin formulation. Pectin-gelatin complexes have been developed [34]; however, there is a distinction between them and our pectin-type B gelatin PECs. First, our pectin-type B gelatin PECs were nano-sized (264.0 ± 3.1 nm), whereas the PECs reported by other research was micro-sized; hence, they are also called complex coacervates. This means that the aggregation that occurred during their manufacture would be different and new, requiring further exploitation. Besides, previous studies mostly reported type A gelatin as the polycation that interacts with polyanionic pectin. The pI of type A gelatin was 9.0, indicating that at pH 7.0, it is positively charged and capable to self-assemble with pectin (pKa 3.0) to form PECs. As pH increased over 9.0, type A gelatin is electrically neutralized. However, the highest pH in the small intestine is approximately 8.0, far lower than the pI of type A gelatin; thus, the drug-releasing site would be mostly only the upper GI tract, such as the stomach. However, D. Madhavi et al. have suggested sustained-release of curcumin via oral administration, which could increase 10-fold its bioavailability compared to that of unformulated curcumin [35]. The drug-releasing efficacy and relative bioavailability of pectin-type A gelatin PEC system would be considered. In this study, type B gelatin was first employed in the PEC system. The pI of type B gelatin was 4.8–5.5 until the pH decreased to 4.0, whereas pectin, whose pKa is 3.0, would be self-aggregated with type B gelatin. At pH 2.0, such as in the stomach, the pectin-type B gelatin complex would swell, as shown in Figure 2, because of the electrically neutralized pectin. The positively charged zeta potential is shown in Figure 3. As the pH increased to 8.0, type B gelatin was electrically neutralized and the zeta potential of the PECs was negatively charged, breaking the electrical balance of the PEC (Figure 3). Increased particle sizes and PDI were therefore observed (Figure 2). The accumulative drug-releasing profile of curcumin-loaded pectin-type B gelatin PECs also revealed their continuous drug-releasing behavior, as shown in Figure 5. Curcumin was not only abruptly released at pH 2.0 but also released sustainably at pH 8.0, which has not been commonly discussed in other studies. The releasing behavior of our PECs differed from that of pectin-type A gelatin PECs. I. Joseph et al. revealed that micro-sized pectin-type A gelatin PECs release their payloads following sustained release at pH 2.0 [36], whereas our PECs rapidly released curcumin at pH 2.0 within 1 h. This finding also evidenced the differences in drug-releasing profiles with different components of the PECs.

Micro-sized PECs (complex coacervates) could bear higher drug contents and loading efficacy after optimization, compared to our PECs. M. Saravanan et al. have prepared pectin-gelatin complex coacervation with particle size of 94 μm and high drug content and loading efficiency, which are 57% and 76%, respectively [26]. Pectin-gelatin complex coacervates were also fabricated by D. Silva el at., with drug content and loading efficiency of 56.7% and 89.5%, respectively [28]. In this study, the drug content and loading efficiency of pectin-type B gelatin PEC were also discussed. The optimum drug content and loading efficiency were 40% and 53%, respectively, whereas the particle size after encapsulation was 214.67 ± 3.43 nm with narrow particular distribution (PDI was 0.17), as presented in Figure 4. The low drug contents and loading efficiency were due to the small cavity of nano-sized PECs. The curcumin-loaded nano-sized PECs prepared by P. Sarika et al. also had low drug content and loading efficiency, compared to those of complex coacervates [24]. The particle size of our curcumin-loaded PECs was slightly smaller than that of bare PEC. The reasonable explanation would be the additional ethanol in the manufacture of these loaded PECs. The ethanol dissolved the curcumin, instead of dispersing it before loading [37]. Thus, the curcumin would not rapidly aggregate inside the core of PECs during manufacture, resulting in similar particle sizes of PEC after curcumin encapsulation.

The drug-releasing profiles of PECs in mimetic GI tract conditions were also investigated in our study, as shown in Figure 5. Curcumin-loaded PECs exhibited abrupt release of curcumin at pH 2.0 and the releasing curve was close to the first-order releasing profiles.  I. Joseph et al. have suggested that the constituents of PECs considerably influence their drug-releasing profile [36]. The drug-releasing profile of our PECs corresponded to the increase in particle size of the bare PECs containing the same ratios of components, as presented in Figure 2a. Particle sizes were observed to abruptly increase as the pH was adjusted to 2.0, which could be attributed to rapid drug-releasing profile at pH 2.0; the particle size became gradually larger at pH 8.0, indicating the sustained-release behavior of the curcumin-loaded PECs. Moreover, the curcumin-loaded PECs preserved almost 40% of payloads within the PECs, which could directly treat colon cancer cells in the GI tract, and the released curcumin would enter the blood via the GI tract to achieve anti-cancer effect. The remaining curcumin-loaded PECs after passing through the mimetic GI tract conditions have been identified that they could be internalized into the colon cancerous cells and further perform the intracellular drug releasing, shown in Figure 6.

The cytotoxicity of the curcumin-loaded PECs to HCT116 human cancer cells was accessed by MTT assay, as presented in Figure 7. In this study, we utilized 1% of DMSO to prepare curcumin solution, which could eliminate the cytotoxicity of DMSO itself [37]. The cell viabilities shown in Figure 7 have indicated that curcumin-loaded PECs exhibited stronger toxicity to HCT116 human colon cancer cells than free curcumin after treatment with mimetic GI tract conditions. The cytotoxicity might be attributed to the anti-cancer activity of curcumin. Y. Wang et al. have shown that active curcumin is degraded in a pH-dependent manner; thus, the alkaline (pH 8.0) treatment might reduce the anti-cancer activity of curcumin [20]. Curcumin-loaded PECs allowed sustained release of curcumin during digestion to maintain curcumin concentration in the blood. Besides, the encapsulated curcumin was protected from external conditions; therefore, the curcumin-loaded PECs could directly carry the bioactive curcumin to cancerous lesions in the colon and efficiently destroyed carcinomas, as evidenced by the results of the internalizing experiments (Figure 6). The cell viability results (Figure 7) indicated that the curcumin-loaded PECs exhibited more significant cytotoxicity to HCT116 cells than free curcumin. The results of cell cycle distribution test (Figure 8) further explained the reason. The cell cycle test results showed that curcumin-treated with mimetic GI tract conditions no longer caused any cell cycle arrest, whereas curcumin-loaded PECs triggered the G2/M phase arrest in HCT116 cells, supporting the results of the studies conducted by A. Jaiswal et al. [38] and J. Lu et al. [39]. Our results have greatly supported that our PECs efficiently preserved the therapeutic effect of curcumin as a payload until it reached carcinomas in the colon.

Curcumin-loaded PECs possessed sustained-release properties and could transport active curcumin into tumorous lesions in the colon after treatment with mimetic GI tract conditions. The results of cell viability tests also confirmed the anti-cancer efficacy of the curcumin-loaded PECs. However, the mimetic GI tract conditions in our study did not involve enzymatic effects, and those conditions were also far from simulating both fasted state gastric conditions (FaSSGF) and fasted state conditions in the small intestine (FaSSIF), as recorded in the pharmacopeia [40]. Future application of curcumin-loaded PECs as a new formulation of oral curcumin solution is still being investigated further but worth expecting.

#### **4. Materials and Methods**

#### *4.1. Materials*

Type B gelatin (molecular weight: 5 k; bloom number: 225), sodium phosphotungstate (PTA) for TEM negative staining and (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT) powder for cell viability assay were all purchased from Sigma-Aldrich Co., Ltd. (St. Louis, MO, USA). Pectin (methylation ratio: 3%–7%) was obtained from Wako Pure Chemical Industries, Ltd. (Osaka, Japan). Curcumin was acquired from Merck Life Science (Hohenbrunn, Germany). Dialysis bags (M.W.C.O.: 6–8 k) for drug-release assessment were also purchased from Merck Life Science. Ethanol was purchased from Taiwan Sugar Corp. (Tainan, Taiwan). Methanol for drug-loading test was obtained from Alps Chem Co., Ltd. (Hsinchu, Taiwan). Reagents for pH value adjustment, hydrochloric acid and sodium hydroxide, were purchased from J. T. Baker Inc. (Center valley, PA, USA) and Uniregion Bio-tech Inc. (New Taipei, Taiwan), respectively. Biological agents, including dimethyl sulfoxide for preparing curcumin solution in cell viability tests and a Muse Cell Cycle Kit was purchased from Merck Millipore (Burlington, MA, USA). Fluorescent reagents, including Cy 5.5-NHS ester, Lysotracker DND-Red, and DAPI-containing mounting medium, were acquired from Lumiprobe (Hunt Valley, MD, USA), Thermo Fisher Scientific (Waltham, MA, USA), and Vector Laboratory (Burlingame, CA, USA), respectively.

#### *4.2. Preparation and Characterization of Pectin-Type B Gelatin PECs*

Pectin (10 mg) was weighed and dissolved in 50 mL of deionized water. Type B gelatin (50 mg) was also weighed and dissolved in 50 mL of deionized water at 25 ◦C. After complete dissolution, the pH of the produced type B gelatin was adjusted to 4.0 using 0.1 N HCl. Ethanol (0.1 mL) was mixed with type B gelatin solution under stirring for 30 min. Next, the pectin solution was added into the mixture, blended, and the pH of the mixture was further adjusted to 4.0. Thirty minutes later, PECs were acquired and stored at 4 ◦C. To optimize the preparation and component of PECs, various volumes of type B gelatin and pectin were added to form the PECs. The type B gelatin/pectin PECs composed of various amounts of pectin and type B gelatin were collected and measured by DLS (Malvern ZS 90, Malvern Instruments, Malvern, UK) for particle size, distribution, and superficial zeta potential (ξ-potential). The hydrodynamic diameter was determined by the Stokes-Einstein equation after DLS measurement. Size distribution was determined through cumulant analysis of the measured intensity autocorrelation function assumed by Gaussian distribution. Zeta potential was obtained through ELS and calculated with Henry's equation.

#### *4.3. Observation of the Morphology of PECs*

The prepared PEC sample was dropped to carbon-coated copper mesh and excess sample was removed after the PEC sample was attached to the mesh. Sodium phosphotungstate aqueous (PTA) solution 1% (*w*/*v*) was prepared and dropped to the copper grids for negative staining. Excess PTA solution was then eliminated. Next, the copper grids were placed in a vacuum oven. Twenty-four hours later, the dried copper grids were kept from light exposure at room temperature until observation. The morphology of the PECs was observed using TEM (HT7700, Hitachi, Japan) with an accelerated voltage of 120 V under high contrast mode (HC mode).

Besides, the morphology of PECs was also observed using a thermal field emission scanning electron microscope (Thermal FE-SEM) (JEOL JSM-7800F, JEOL Ltd., Tokyo, Japan). The PEC solutions were first prepared and freeze-dried using a vacuum freeze drier (Uniss MF-280, Uniss Company, New Taipei City, Taiwan). The freeze-dried powder was afterwards coated with platinum and observed by the thermal FE-SEM.

#### *4.4. pH Responsiveness of PECs*

PECs (5 mL) were placed into a sample vial and its pH was adjusted to 2.0 using HCl (0.1 N) to mimic the acidic condition in the stomach. Afterward, the PEC solution was incubated at 37 ◦C for 2 h in a shaker (M.BR-022UP, Taitec, Tokyo, Japan). Next, the pH of the PEC solution was adjusted to 8.0 using NaOH (0.5 N) to simulate the alkaline environment in the small intestine. The adjusted PEC solution was placed at 37 ◦C for 4 h in a shaker. Particle size and surface zeta potential were measured by DLS at every hour.

#### *4.5. Preparation and Characterization of Curcumin-Loaded Pectin-Type B Gelatin PECs*

Different amounts of curcumin (0.5, 1, 1.5, 2, or 2.5 mg) were precisely weighed and placed independently in a vial. Curcumin was first dissolved in ethanol (0.1 mL). Simultaneously, pectin (10 mg) and type B gelatin (50 mg) were respectively weighed and dissolved in 50 mL deionized water as stock solutions. Type B gelatin solution was treated with 0.1 N of HCl until the pH was 4.0. Curcumin alcoholic solution (0.1 mL) and type B gelatin solution (0.2 mL) were first mixed in a sample vial with stirring for 30 min. Thereafter, pectin solution (0.7 mL) was added to the mixture, which was then blended for 30 min to fabricate curcumin-loaded PECs. Particle size and distribution were determined using DLS.

#### *4.6. Drug Contents and Loading Efficiency Determination*

Curcumin-loaded PECs were prepared as abovementioned. The prepared curcumin-loaded PECs were freeze-dried using a vacuum freeze drier (Uniss MF-280). The dried sample (0.5 mg) was then weighed and dissolved in 1 mL of methanol for curcumin extraction. The extracted curcumin was quantified by an ELISA reader (BioTek Synergy HT, BioTek Instruments, Inc., Winooski, VT, USA) at 423 nm as the detected wavelength.

The drug content of the curcumin-loaded PCs was calculated according to the formula (1):

$$\text{Drug contents } (\%) = \frac{weight \, of \, encapsulated \, curcumin \, (mg)}{total \, weight \, of \, curcumin - loaded \, PECs \, (mg)} \times 100\% \tag{1}$$

Besides, drug loading efficiency was also determined and calculated according to the formula (2):

$$\text{Loading efficiency } (\%) = \frac{encapsulated \, curcumin \, weight \, (mg)}{total \, circumin \, input \, weight \, (mg)} \times 100\% \tag{2}$$

After calculation, the optimization was analyzed with analysis of variance (ANOVA).

#### *4.7. Determination of Drug-Releasing Profiles*

Curcumin-loaded PECs (1 mL) were placed into dialysis bags, which were then immersed into 2.5 mL of methanol/deionized water solution (1:1, *v*/*v*) in a sample vial. The pH of the methanol/deionized water solutions in a sample vial was first adjusted to 2.0, and then the sample vial was incubated at 37 ◦C for 2 h. Afterwards, the pH value of the methanol/deionized water solution in the sample vial was adjusted to 8.0, and the sample vial was further incubated at 37 ◦C for 4 h. At each hour, 0.1 mL of methanol/deionized water solution was collected and absorbance wavelength at 423 nm was detected to quantify the release of curcumin.

#### *4.8. CLSM Observation*

HCT116 cells (1 × 105 cells/mL) were seeded on a slide and incubated at 37 ◦C with 5% CO2 supply for 12 h. After the cells were attached, they were co-cultured with free curcumin- or Cy5.5-labeled curcumin-loaded PECs, which were pretreated with mimetic GI tract conditions for 12 h at 37 ◦C with 5% CO2 supply. Excess curcumin or curcumin-loaded PECs were thereafter removed from the cells, which were then washed twice with PBS. The cells were further stained with 1 μM of Lysotracker DND-Red at 37 ◦C with 5% CO2 for 1 h. Next, the lysotracker DND-Red was removed, and the cells were again washed twice with PBS and fixed with 4% paraformaldehyde for 20 min. After fixation, the cells were washed twice with PBS and the cells and their nuclei were stained and mounted with DAPI-containing mounting medium. Fluorescence was observed and detected using a confocal laser scanning microscope (Leica SP8, Leica Microsystem, Wetzlar, Germany). The fluorescence of Cy 5.5 was detected at excitation and emission wavelengths of 633 and 650 nm, respectively. The fluorescence of Lysotracker DND-Red for locating endosome/lysosomes was detected using excitation and emission wavelengths of 577 and 590 nm, respectively. The fluorescence of curcumin was observed at excitation and emission wavelengths of 488 and 520 nm, respectively. DAPI fluorescence was detected using the preset wavelength of the CLSM equipment.

#### *4.9. Cell Viability Determination*

Human colon cancer cells HCT116 (7 × 103 cells/mL) were seeded on a 96-well plate at 37 ◦C with 5% CO2 supply for 12 h to allow cell attachment. Simultaneously, curcumin-loaded PECs were adjusted with the cell culture medium to the predetermined concentrations at 45 to 1.41 μM. Curcumin solutions at the same concentration series were also prepared with 1% of DMSO and cell culture medium. Afterwards, the pH of both curcumin and curcumin-loaded PECs was adjusted to 2.0 for 2 h, then the pH of the PECs was adjusted to 8.0 for 4 h. The pH of the PECs was re-adjusted to 7.4, and then HCT116 cells were treated with the PECs. The HCT116 cells were incubated at 37 ◦C with 5% CO2. Twelve hours later, the PECs and cell culture medium were removed. MTT assay was utilized to evaluate cell viability.

In addition, the human normal fibroblast cells Detroit 551 (7 × 103 cells/mL) were also seeded on a 96-well plate and cultured at 37 ◦C with 5% CO2 for 12 h. Bare PECs at concentrations of 112.5 to 3.52 μM were prepared with serial dilution. The bare PECs were separated into two groups: one for direct treatment with Detroit 551 cells and the other for mimetic GI tract treatment, as described above. Next, the latter bare PECs were transferred to pH 7.4 condition, and the cells were cultured with the bare PECs for 12 h. MTT assay was then applied to determine cell viability.

#### *4.10. Cell Cycle Analysis*

HCT116 cells (3 × 106 cells/mL) were seeded on a 6-well plate and cultured for 12 h at 37 ◦C with 5% CO2 supply. After the cells were attached, they were exposed to curcumin or curcumin-loaded PECs pretreated in mimetic GI tract conditions (pH 2.0 for 2 h, pH 8.0 for 4 h, and back to pH 7.0), as described above, and then the cells were incubated at 37 ◦C with 5% CO2 supply for 12 h. Afterward, the cells were collected, washed with PBS, and fixed with 70% of ethanol at −20 ◦C for 6 h. Fixed cells were re-collected using centrifugation and suspended in PBS. The cells were further stained with a flow cell cycle kit in the dark for 30 min. Cell cycles were determined using flow cytometry (BD FACSCanto, BD Biosciences, Franklin Lakes, NJ, USA) and analyzed with the software ModFit LT 3.0 (BD Biosciences, Franklin Lakes, NJ, USA).

#### *4.11. Statistical Analysis*

All experiments were repeated at least thrice and the results were averaged and presented as mean ± standard deviation (S.D.). Most results have been statistically compared and analyzed using Student's *t*-test (Microsoft Excel 2000). Differences were recognized as statistically significant when the *p* values were less than 0.05, and significant differences are shown in star marks (\* *p* < 0.05; \*\* *p* < 0.01, and \*\*\* *p* < 0.001). Besides that, of all results, only the optimization of drug contents and loading efficiency were analyzed with analysis of variance (ANOVA) using OriginPro 8 (OriginLab, Northampton, MA, USA).

#### **5. Conclusions**

In this study, a novel pectin-type B gelatin PEC was fabricated for curcumin delivery and its optimum manufacturing processes, including component ratios and curcumin-loading process, were also developed. The produced PECs could increase their particle size at pH 2.0 for the neutral charges of pectin. At pH 8.0, type B gelatin would be electrically neutralized, driving the gradual swelling of the PECs. The curcumin-loaded PECs could release their payloads at both pH 2.0 and 8.0, which potentiate exploration of their oral formulation. Besides, the PECs could also preserve the bioactivity of curcumin and be internalized into colon cancer cells to allow intracellular drug-releasing behaviors. In addition, the curcumin-loaded PECs induced cell cycle arrest and excellent cytotoxicity toward the human colon cancer cells HCT116 even after treatment under mimetic GI tract conditions, whereas bare PECs by themselves showed very low toxicity to normal cells. The pectin-type B gelatin PEC that possessed high safety and outstanding anti-cancer effects is a promising oral formulation for curcumin delivery.

#### **Supplementary Materials:** Supplementary materials can be found at http://www.mdpi.com/1422-0067/19/11/ 3625/s1.

**Author Contributions:** Y.-T.C. and I.-J.S. conceived and designed the concepts and experiments. Y.-T.C. and X.L. analyzed the results and wrote the manuscript. F.-Y.S. and L.-L.C. performed the preparation and characterization of bare PECs, including particle size and pH responsiveness. F-Y.S. also optimized the drug loading and conducted drug releasing alone. F.-Y.S. conducted the cell culture and the relative experiments, such as MTT assay and cell cycle arrest analysis under the supervision by S.-C.K. and Y.-C.H.

**Funding:** This research was funded by Ministry of Science and Technology, Taiwan (MOST) (MOST 106-2320-B-039 -001 and 107-2314-B-039-013) and by China Medical University (CMU106-N-10 and CMU107-N-08). The research was also financially supported by the "Chinese Medicine Research Center, China Medical University" from the Featured Areas Research Center Program within the framework of the Higher Education Sprout Project by the Ministry of Education (MOE) in Taiwan (CMRC-CHM-6).

**Acknowledgments:** Particle size analysis was performed through the use of the Medical Research Core Facilities Center, Office of Research & Development at China medical University, Taichung, Taiwan. TEM observation was conducted by the use of the Instrument Center at National Tsing Hua University, Hsinchu, Taiwan. The authors would especially thank Ju-Chun Tai for several equipment operations, including CLSM and flow cytometry at Medical Research Core Facilities Center, Office of Research & Development at China medical University, Taichung, Taiwan.

**Conflicts of Interest:** The authors declare no conflict of interest.

#### **Abbreviations**

#### **References**

© 2018 by the authors. Licensee MDPI, Basel, Switzerland. This article is an open access article distributed under the terms and conditions of the Creative Commons Attribution (CC BY) license (http://creativecommons.org/licenses/by/4.0/).

## *Article* **Chemotherapeutic Efficacy of Implantable Antineoplastic-Treatment Protocols in an Optimal Mouse Model for Human Ovarian Carcinoma Cell Targeting**

**Jonathan M. Pantshwa 1, Khadija Rhoda 1, Sarah J. Clift 2, Priyamvada Pradeep 1, Yahya E. Choonara 1, Pradeep Kumar 1, Lisa C. du Toit 1, Clement Penny 3 and Viness Pillay 1,\***

Received: 23 June 2018; Accepted: 27 September 2018; Published: 4 October 2018

**Abstract:** The present study aimed to design and develop a nanocomposite drug delivery system employing an antineoplastic-loaded antibody functionalized nanomicelle encapsulated within a Chitosan–Poly(vinylpyrrolidone)–Poly(*N*-isopropylacrylamide) (C–P–N) hydrogel to form an in situ forming implant (ISFI), responsive to temperature and pH for cancer cell-targeting following intraperitoneal implantation. The optimum nanomicelle formulation was surface-functionalized with anti-MUC 16 (antibody) for the targeted delivery of methotrexate to human ovarian carcinoma (NIH:OVCAR-5) cells in Athymic nude mice that expressed MUC16, as a preferential form of intraperitoneal ovarian cancer (OC) chemotherapy. The cross-linked interpenetrating C–P–N hydrogel was synthesized for the preparation of an in situ-forming implant (ISFI). Subsequently, the ISFI was fabricated by encapsulating a nanocomposite comprising of anti-MUC16 (antibody) functionalized methotrexate (MTX)-loaded poly(*N*-isopropylacrylamide)-b-poly(aspartic acid) (PNIPAAm-b-PASP) nanomicelles (AF(MTX)NM's) within the cross-linked C–P–N hydrogel. This strategy enabled specificity and increased the residence time of the nanomicelles at tumor sites over a period exceeding one month, enhancing uptake of drugs and preventing recurrence and chemo-resistance. Chemotherapeutic efficacy was tested on the optimal ovarian tumor-bearing Athymic nude mouse model and the results demonstrated tumor regression including reduction in mouse weight and tumor size, as well as a significant (*p* < 0.05) reduction in mucin 16 levels in plasma and ascitic fluid, and improved survival of mice after treatment with the experimental anti-MUC16/CA125 antibody-bound nanotherapeutic implant drug delivery system (ISFI) (*p* < 0.05). The study also concluded that ISFI could potentially be considered an important immuno-chemotherapeutic agent that could be employed in human clinical trials of advanced, and/or recurring, metastatic epithelial ovarian cancer (EOC). The development of this ISFI may circumvent the treatment flaws experienced with conventional systemic therapies, effectively manage recurrent disease and ultimately prolong disease-free intervals in ovarian cancer patients.

**Keywords:** implant; antibody functionalized nanomicelles; epithelial ovarian cancer; chemotherapeutic drugs; nude mouse model

#### **1. Introduction**

Ovarian cancer is one of the fatal female diseases in which many patients are diagnosed at an advanced stage when the disease has already spread beyond the ovaries to the abdominal cavity [1–3]. Targeted treatment using biomarkers for OC holds great promise to improve the survival rate of patients with OC. Actively targeted chemotherapy has been a considerable advancement compared to passive-targeting based on the Enhanced Permeability and Retention (EPR) effect. To this end, actively targeted chemotherapeutic drug delivery employing nanotechnology has had a remarkable impact on cancer treatment with the following advantages: (1) Ability to deliver antineoplastic drugs to treat specific cancer metastatic sites; (2) potential reduction in the quantity of drug necessary to achieve a required therapeutic dose at the target cancer tissue; and (3) reduction in the quantity of drug delivered to healthy cells thus decreasing cell cytotoxicity [4–6]. In this context, the coating of nanomicelles with specific antibody targeting ligands can assist in conjugation of the nanomicelles to precise biomarkers and antigen receptors over-expressed on OC cells to target tumors with improved accuracy [7,8]. The flexibility of coated amphiphilic nanomicelles is amenable to modified forms of chemotherapy. Targeting cancer cells at a molecular level using nanocarriers such as nanomicelles may be the solution for more personalized chemotherapy. Antibodies are the most significant target ligands that have offered a broad-spectrum of possibilities in terms of drug targeting and accuracy of interaction. Antibody-targeted nanomicelles (or immunomicelles) can be engineered by composite conjugation of antibodies (or nanobodies) onto the trigger (exposed free terminal) of a hydrophilic segment of the nanomicelle. This ensures active targeting nanomicelle preparation without any steric-hindrance for the ligand-antibody [9,10].

Therefore, this study has innovatively focused on the use of mucins (specifically MUC16) for the advanced design of a novel intraperitoneally administered (as opposed to conventional intravenous chemotherapy) OC cell-targeting nanocomposite drug delivery system to significantly improve the chemotherapy of OC. The uniqueness of the delivery system is via the use of antibody-functionalized nanomicelles to specifically target mucin antigens known to be over-expressed in OC cells. To our knowledge, this has not been investigated before. MUC16, specifically, is an eminent cell surface antigen in OC that is shed into the serum and therefore it is also widely used clinically for the diagnosis and management of epithelial OC [11–13].

The monoclonal antibody OC125/Mab anti-MUC16 identifies this tumor-associated antigen, which is over-expressed in epithelial ovarian cancer cells, but not in the epithelium of normal fetal and mature ovaries [14]. The non-specific delivery of chemotherapeutics to healthy tissues other than the OC tumors is one of the leading cytotoxicity challenges of chemotherapy. Therefore, this study also focuses on surmounting this challenge by designing MUC16 antibody-functionalized antineoplastic drug-loaded nanomicelles to specifically target OC cells via the peritoneum. However, targeted nanomicelles have a few limitations regarding the ability to provide sustained drug release. The limitations of these nano-enabled polymeric systems include instability in the circulatory system, rapid degradation, and clearance by the immune system and a lack of controlled sustained drug release [15,16]. To overcome these limitations for clinical applications to be feasible, current research focuses on preparation of nanomicelles encapsulated with hydrogels for solid intra-tumoral release of chemotherapeutic drugs over-prolonged periods of weeks, or even months, thus diminishing the dire need for daily administration of systemic chemotherapy. Hence, prolonged circulation in the bloodstream, in-vivo stability, biodegradability, and polymer-drug compatibility with sufficient retention of the drug within the carrier system are prerequisites to successful design and preparation of an implantable drug targeting delivery system [17,18].

#### **2. Results**

#### *2.1. Mechanism of Synthesis of the Anti-MUC16 MTX-Loaded Nanomicelles*

Amphiphilic block copolymers are expected to offer improved applications in chemotherapeutics when formulated as nanomicelles functioning as nanocarriers for ionic and non-ionic drug loading [19–21]. Ionic complexes/composites involving anionic antineoplastic drugs and aspartic acid can be produced and nanomicelles further synthesized by self-assembly of the drug and amphiphilic copolymer in an aqueous medium. In this amphiphilic structure, the hydrophilic surface shields the hydrophobic interior from the interaction of the inner core with the surrounding aqueous environment, thus enabling the nanomicelles to be used as a vehicle for targeted drug delivery (Figure 1).

**Figure 1.** A schematic illustration of the synthesis of an implantable antibody functionalized methotrexate loaded nanomicelles hydrogel composite delivery system (ISFI). Blue circles refer to nanomicelle's corona.

In this study, a more efficient strategy (under moderate formulation conditions) was utilized to ensure superior targeting of the functionalized core-shell nanomicelles using a simplified synthetic approach outlined in Figure 1. The physical synthesis comprised two synthetic steps: (1) The block copolymer constituted an aspartic (ASP) backbone and poly(*N*-isopropylacrylamide) (PNIPAAm) outer surface to self-assemble the MTX-loaded core-shell nanomicelles by solvent evaporation in an aqueous medium. The polycationic backbone formed between poly(aspartic acid) (PASP) and methotrexate (MTX) generated the inner core of the nanomicelle, while the non-ionic PNIPAAm formed the outer hydrophilic shell. PASP was used to conjugate the MTX (a carboxylic drug) to its amine groups to form an amide (CO-NH) linkage (2). Furthermore, the anti-MUC16 antibody was surface-coated to the nanomicelles via carbodiimide-sulpho-NHS mediated conjugation through carboxylic (-COOH) linkage of the PNIPAAM-b-PASP copolymer. This linkage was also significant for the elucidation of the orientation of the anti-MUC16 to ensure binding affinity and stability. In this case, the MTX-loaded nanomicelles were incubated with 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide and *N*-hydroxysuccinimide for 15 min at room temperature. The resulting activated nanomicelles were then covalently linked to the anti-MUC16 antibody (1% weight compared with the polymer concentration) (Figure 1). The resultant anti-MUC16 antibody-functionalized MTX-loaded nanomicelles (the NanoComposite) were then lyophilized (Freezone 6, Model 79340, Labconco, MO, USA).

The results of this study confirmed that MTX can be loaded into the interior of the PNIPAAM-block-PASP nanocomposites and the carboxyl-amine-groups of the nanomicelles were amenable to surface coating with the anti-MUC16 antibody (Figure 1). It is well known that MUC16 is highly expressed on malignant cells and not on healthy cells and therefore the NanoComposite can impede tumor growth and even halt metastasis based on the binding to MUC16. In addition, it is important to note that OC cells also use MUC16 as a mechanism for metastasis via muco-adhesion to other organs within the peritoneum. Thus, the developed NanoComposite has far-reaching potential for the targeted treatment and management of OC employing advanced nanochemotherapeutics.

#### 2.1.1. Analysis of the Molecular Structure Integrity of the Functionalized Nanocomposite

Fourier-transform infrared (FTIR) spectral analysis is one of the most robust analytical techniques to evaluate the molecular structure of block copolymer functional groups [22,23]. The interactions and functionalization of the MTX-loaded nanomicelles with anti-MUC16 was evaluated and described with FTIR spectral analysis and is presented in Figure 2.

As demonstrated, there were distinct chemical structure transitions for the functionalized MTX-loaded nanomicelles compared with the non-functionalized MTX-loaded nanomicelles. The FTIR spectra of the blank nanomicelles and MTX-loaded NanoComposite were congruent with those of the native polymers PNIPAAm and ASP [24,25]. This result showed that the polymer formation into nanomicelles occurred with minor chemical modification during the polymerization process. Hence, the nanomicelles possessed elemental chemical features that were representative of the native polymers.

Variations were noted in the FTIR spectrum of the blank nanomicelles and MTX-loaded nanomicelles in Figure 2. The additional peaks that were recorded in the MTX-loaded nanomicelles were due to the 1,3-replacement composite (1516.53–1451.23 cm−1) and a phenyl amino compound (1647.22–1451.13 cm−1). This indicated that MTX was entrapped within the hydrophobic interior of the nanomicelle structure by weak electrostatic H-bonds involving the COO- moiety of MTX and the OH-group of aspartic acid or by the charged ionic groups involving the NH2 bonds of MTX and the COO-bonds existing in PNIPAAm-b-PASP.

MTX was dispersed within the PNIPAAm-b-PASP matrix in the amorphous form exclusive of any polymorphic transformation or alteration in the aqueous form. The anti-MUC16 amine groups (N–H) were evident by the 1660.20 cm−1 peak as well as the derived amine groups induced by the broad range O–H group and the C–H group observed between 1090.30–1310.11 cm−1 that also overlapped the 1000.12–1200.22 cm−1 band representative of C–O groups that were attributed to the PNIPAAm-b-PASP copolymer. These transitions indicated that there were prominent interactions

between the MTX-loaded nanomicelles and the anti-MUC16 antibody that confirmed the surface coated functionalization of the MTX-loaded NanoComposite.

**Figure 2.** Fourier-transform infrared (FTIR) spectra illustrating the molecular structural transitions of MTX, blank micelles, AF(MTX)NM's and (MTX)NM's.

2.1.2. Analysis of the Thermal Features of the Drug-Free and MTX-Loaded PNIPAAm-b-PASP Nanomicelles

Differential scanning calorimetric investigations were conducted to observe the drug's physical form in the nanomicelle, since this could affect the drug release from the delivery system. DSC thermograms of MTX, blank micelles and MTX-loaded micelles are depicted in Supplementary Data Figure S1. The DSC curve of the MTX displayed a critical melting peak at 125 ◦C. MTX-loaded micelles displayed the melting peak for the drug at 145 ◦C. Incorporation of MTX within the nanomicelle thus elevated the melting point indicative of enhanced thermal stability of the drug. This would suggest that on incorporation of the partially soluble crystalline MTX within the hydrophobic core it was converted to a less crystalline form, resulting in a favorable MTX entrapment efficacy of 80.6 ± 0.3%. The glass transition temperature (*Tg*) of the co-polymeric nanomicelles is also evident at ~150 ◦C, correlating with previous findings [25].

#### *2.2. Synthesis and Characterization of Chitosan-Poly(N-Vinylpyrrolidone)-Poly(N-Isopropylacrylamide) (C–P–N) Hydrogel*

Figure 3 illustrates the constituent chemical structures and mechanism of synthesis of the C–P–N composite hydrogel structure. Physicochemical characterization via nuclear magnetic resonance (NMR) and FTIR analysis indicated the formation of chitosan-poly(*N*-vinylpyrrolidone) poly(*N*-isopropylacrylamide) covalent bonds between the protonated amine groups of chitosan (annotated with red circles) and carboxylated groups of poly(*N*-vinylpyrrolidone) (annotated with green circles). Both covalent crosslinking and chemical structure moderated the swelling of the IPN hydrogel. The synthesized C–P–N hydrogel was also verified with chemical structural correlation and proton NMR crest peak assignments (Supplementary Materials Figure S2).

**Figure 3.** Illustration of constituent chemical structures and reaction mechanism forming composite hydrogel structure.

Assessment of Polymeric Structural Variations of the Cross-Linked C–P–N Hydrogel

FTIR analysis of the C–P–N hydrogel proved successful radical polymerization of chitosan (CHT), polyvinylpyrrolidone (PVP) and PNIPAAm, and interpenetration of glutaraldehyde (GA) within the cross-linked network (Supplementary Materials Figure S3). The FTIR spectrum of the C–P–N hydrogel demonstrated peaks at 3428 and 1654 cm−1 assigned to stretching vibrations of -NH and -OH, respectively. The peak at 1720 cm−1 was attributed to the stretching vibrations of carbonyl group in the PVP-PNIPAAm molecule which differentiated chitosan and the chitosan derivative (CHT-PVP). The PNIPAAm hydrogel sample exhibited significant peaks at 1654, 1551, 1385 and 1369 cm−1, which were assigned to the characteristic peaks of amide I, amide II and the isopropyl group, respectively (Supplementary Materials Figure S3). For the composite hydrogel, the intensity of the absorption peak at 1654 cm−1 was attributed to increased amide group compared with the CHT-PVP hydrogel due the incorporated NIPAAm. The intensity of the methyl peak and isopropyl peaks decreased when compared with PNIPAAm hydrogel. This indicated successful radical polymerization of CHT-PVP-PNIPAAm and interpenetration by GA and *N*,*N*-methylenebisacrylamide.

#### *2.3. Characterization of the In Situ Forming Implant (ISFI)*

2.3.1. Assessment of the Gelation Temperature Using Oscillatory Rheology

The crossover of the storage and loss modulus (G and G") indicates the gelation temperature (*Tgel*) of the hydrogel as illustrated in Figure 4. The storage modulus (G ) of a viscoelastic solid is associated with the solid properties or the elastic energy storage properties which indicated that the sample would return to its original state following removal of the deformational energy, while the loss modulus (G") explains the behavior of the viscoelastic solid when it is acting as a liquid i.e., the viscous properties of the sample representing the dissipation of energy after the application of the deformational stress. Hence at the point at which the storage modulus exceeded the loss modulus, the C–P–N hydrogel was behaving more like a solid than a liquid and hence this was used to determine the thermal gelation temperature. Gelation studies were also conducted on non-crosslinked samples and used as controls (Figure 4).

**Figure 4.** Typical profile obtained during oscillatory rheological testing of non-crosslinked (**a**) and crosslinked (**b**) C–P–N hydrogel samples. The black and blue circles demarcates the crossover point storage modulus (G ) and the loss modulus (G").

#### 2.3.2. Morphological Characterization of the ISFI

The structure and properties of the ISFI were validated by confocal laser scanning microscopy (CLSM) and scanning electron microscopy (SEM). Figure 5 markedly illustrates the morphological properties of the drug-filled coated nanocomposites after encapsulation in C–P–N gel of the ISFI. Drug-loaded coated nanocomposites have a homogenous spherical form. CLSM confirmed optical distribution of the labeled coated nanocomposites in the spatial polymeric-grounded depot platforms. Control native or unlabeled C–P–N hydrogel possessed no rhodamine activities. The overall results

substantiate that drug-loaded coated nanocomposites maintain their integrity after-lyophilization for 48 h.

**Figure 5.** Typical SEM micrographs of the drug-loaded coated nanomicelles after encapsulation in the C–P–N at 50× lower magnification (**a**,**c**); and (**b**) 100× Increased magnification of the ISFI.

To fully understand the potential of utilizing the ISFI for local drug delivery to human ovarian carcinoma cell, the first factor that had to be assessed was the cytotoxicity. Investigation of in vitro cytotoxicity was conducted on the nanomicelles via MTT assay of OVCAR 5 cells in our previously published work [26]. Similar trends were observed in this study, except that the ISFI enabled a prolonged drug release of 30 days.

#### *2.4. In Vivo Studies in the Athymic-Nude Mice Model*

In vivo studies were conducted in three platforms: Firstly intraperitoneal (IP) induction of human ovarian carcinoma in Athymic Swiss nude mice, secondly IP implantation of various treatment protocols including ISFI on the optimal ovarian tumor-bearing Athymic nude mouse model, and thirdly in vivo evaluation of chemotherapeutic efficacy utilizing a variety of indices including whole mouse weight, tumor size (measured with calipers and sonography), quantification of mucin 16 antigen expression levels (using immunohistochemistry), and sandwich quantitative CA125 Human ELISA (on plasma and ascitic fluid), as well as the survival rate of mice post-treatment.

#### 2.4.1. Induction of Human Ovarian Carcinoma in Athymic Swiss Nude Mice

Growth of cancer cells in the intra-peritoneal cavity and subcutaneous region of immuno-compromised mice is a common procedure for evaluating tumorigenic potential in vivo [19–21]. These procedures are also utilized to evaluate the effects of chemotherapeutic interventions on tumor cell lines. Charles River France Swiss nude mice (*n* = 80), were established as a simple, reproducible mouse model for the induction of human ovarian carcinoma using the IP inoculated NIH:OVCAR5 cell line, as was evidenced by the visible development in vivo of intra-abdominal tumor nodules with associated severe ascites. This occurred within 10 days of inoculation. Advanced ovarian carcinoma disease in the Athymic nude mice was consistently associated with peritoneal carcinomatosis/transcoelomic metastasis, which always preceded the formation of severe ascites. Solid tumor nodules coated all serosal surfaces, especially within the pelvis. These tumor-bearing mice were further IP implanted with various treatment protocols and chemotherapeutic efficacy was evaluated. Due to the widespread dissemination of the IP tumors, it was not feasible to analyze specific regions of the peritoneum to assess the uptake of the antibody (anti-MUC16) and drug-loaded nanomicelle delivery system in the peritoneum.

#### 2.4.2. Chemotherapeutic Efficacy in the Treatment of Human Ovarian Carcinoma

The chemotherapeutic efficacy of the AF(D)NMs implant (the preferred chemotherapeutic model system) was evaluated against the non-functionalized (D)NMs and the comparison group (IV methotrexate/cisplatin only), as well as the control/placebo group. A variety of indices were assessed, including tumor size (measured with calipers and sonography), mouse weight, quantification of mucin 16 antigen expression levels as well as survival rate of mice post-treatment.

#### 2.4.3. Tumor Size

The tumor size was measured with calipers and sonography (Figure 6) and two chemotherapeutic drugs utilized in this study were cisplatin (4 mg/kg) and methotrexate (15 mg/kg), as illustrated with nude mice IP-growth curves, respectively (Figure 7A,B). The average tumor size in the 3 treatment groups each with 10 mice (two experimental groups, and a comparison group (i.v group)) decreased significantly (*p* < 0.05) from day 15 after implantation of the AF(D)NMs treatment. Conversely, in the placebo group, the average tumor size increased steadily, indicating biocompatibility of the blank nanomicelle implant delivery system (placebo) in vivo. During the evaluation phase, nanomicelle implant ((D)NMs)-treated mice reached the ultimate point (of 100 mm3 average tumor diameter) within 21 days of treatment whilst the group of AF(D)NMs-implant treated mice survived until the completion of the study. After necropsy examination of the IP inoculated mouse post treatment, the AF(D)NMs implant treatment resulted in reduced average tumor size and ascitic fluid.

**Figure 6.** Sonographic representation of IP-tumor growth and response to antibody-bound drug loaded nanomicelle hydrogel composite delivery system (AF(D)NMs). (**a**) Ascitic fluid development in a nude mouse 5 days post-induction; (**b**) Tumor growth 10 days post induction with NIH:OVAR-5 cell suspension; (**c**) Chemotherapeutic implant injected IP-adjacent to tumor growth 11 days post-induction, (**d**) decrease in tumor size and only a small tumor nodule was noticeable 15 days after implementation of the IP-(AF(D)NMs) treatment.

**Figure 7.** Nude mouse IP-tumor growth curves illustrating chemotherapeutic efficacy of the three treatment groups vs. the control/placebo group, expressed as average tumor sizes in NIH:OVCAR-5 EOC-bearing nude mice. (**A**) refers to the methotrexate and (**B**) the cisplatin-loaded nanomicelle implant delivery system in NIH:OVCAR-5 EOC-bearing nude mice. Each point depicts mean (*n* = 10/group); bar, ±SD. \* *p* < 0.005 by *t*-test.

#### 2.4.4. Whole Mouse Weight

Mouse weights in both the comparison group (IV cisplatin/methotrexate group) and the 2 experimental treatment groups decreased during this study (*p* < 0.05) (Supplementary Materials Figure S4). Conversely, in the placebo group body weights increased slightly and were normalized to baseline weight (*p* < 0.05). The final average weight of nude mice with NIH:OVCAR-5 tumor treated with AF(D)NMs formulations was 16.94 ± 0.3 g compared with 17.68 ± 0.3 g in mice treated with methotrexate/cisplatin–loaded nanomicelles (D)NMs and 26.34 ± 0.36 g in mice administered with placebo injections (*p* < 0.05). The IV drug group final average weight was 22.62 ± 0.28 g. Baseline weights were 26.00 ± 0.40 g, 24.00 ± 0.33 g and 19.1 ± 0.35 g for placebo, (D)NMs, IV drug and AF(D)NMs, respectively. The body weight of the placebo group was normalized to 26.00 g on

day 30, indicative of implant biocompatibility and low levels of cytotoxicity (Supplementary Materials Figure S4).

The notable reduction in tumor size and mouse weights corresponded with decreased MUC16/CA125 antigen expression levels determined with the ELISA technique. These results indicated that there was significant difference (*p* > 0.01) in tumor burden between different chemo-treatment groups and confirmed that antibody functionalized combinational treatment significantly improved chemo-therapeutic efficacy as demonstrated by the inhibition of tumor growth (*p* < 0.05).

#### 2.4.5. Quantification of Plasma and Ascitic Fluid MUC16/CA125 Antigen Levels

Using the Cancer Antigen CA125 Human ELISA Kit, MUC16 antigen concentrations in IV-treated mouse plasma samples were significantly lower (*p* < 0.05) than the pre-treatment group. However, MUC16 antigen concentrations in mice treated with (D)NMs and AF(D)NMs were all in a lower range, i.e., 1.8–2.4 U/mL. The MUC16 antigen concentration in the plasma samples was, typically, slightly more elevated than in the ascitic fluid, due to the AF(D)NMs site-specificity and localization in the peritoneal cavity. AF(D)NMs target MUC16 antigens expressed on ovarian carcinoma cells within tumor nodules and in ascitic fluid. This AF(D)NMs group significantly reduced MUC16/CA125 antigen concentrations in plasma, as compared to (D)NMs group and IV chemotherapeutic drug group (*p* < 0.05) (Figure 8). At the time of euthanasia, the average plasma MUC16/CA125 concentration in the AF(D)NMs group was 1.9460 U/mL, compared to 2.0180 U/mL in the (D)NMs group, 2.077 U/mL in the IV drug group and 2.368 U/mL in the placebo group. The decrease in MUC16/CA125 antigen concentration in ascitic fluid was consistent with the overall reduction in ascitic fluid production and average tumor size. These results indicate that the AF(D)NMs can specifically target MUC16/CA125 antigens on the surface of EOC cells, thereby effectively decreasing their expression.

**Figure 8.** Point-of-euthenasia levels of MUC16/CA125 antigens in the plasma and ascitic fluid of mice in the experimental and control groups. (D) refers to methotrexate in (**a**) and cisplatin in (**b**). Thick and thin bars refer to pre-treatment and post-treatment groups, respectively. \*: *p* < 0.05 by *t*-test.

#### 2.4.6. IP-Tumor Histopathology

Gross necropsies were performed on all the pre-treatment animals that died or were euthanized by carbon dioxide (CO2) inhalation after 10 days. There was no difference in histomorphological appearance of IP-inoculated tumors. Histopathology was subsequently performed in all IP nodules and in all cases, the existence of anaplastic ovarian carcinoma(s) was confirmed (Figure 9). The intra-abdominal tumors, however, were far more infiltrative (widespread transcoelomic metastases with extensive regions of necrosis and histopathology also revealed tumor emboli within lymphatic vessels). Necrosis was linked with a diffuse inflammatory response, comprising of mononuclear cells (Figure 10b). The inflammatory process on the tumor site is controlled by the neutrophil-infiltration. Macrophages are activated to release TNF-α by neutrophil chemotactic factors and reactive oxygen species resulting in the elimination of tumor cells, however an uncontrolled pattern can promote cell laceration and necrosis. After a month of chemotherapy, the mice in the post-treatment groups were euthanized in the same manner as those in the pretreatment group and full liver histopathology performed.

#### 2.4.7. Liver Histopathology

In contrast, tissues sections of nude mice of the treated group demonstrated minor necrotic regions and a lower number of IP-tumor cells (Figure 9). The reduced necrosis was consistent with the decrease in IP tumor-growth. In this vein, histopathology performed on IP nude mice implanted with experimental IV and placebo treatments displayed multifocally coalescing neoplastic nodules throughout the peritoneal cavity, as well as multiple random foci of hepatocellular coagulative necrosis associated with bile "lakes" (so called "bile infarcts") likely due to biliary outflow obstruction by ovarian carcinomas (in several instances, carcinomatous foci were observed immediately abutting extra-hepatic biliary cysts (Figure 9a,b)). There was also evidence of neoplastic emboli in some sections in the placebo group. There was very mild bile ductile proliferation within the liver sections as well as occasional extra-hepatic biliary cysts (lined by hyperplastic epithelium), the latter also embedded within an increased fibrous connective tissue stroma. Numerous bile ductules within the liver were bile-laden and some bile ductules were severely distended in portal areas and lined by hyperplastic (in certain regions, pseudoepitheliomatous) epithelium. There was a mild portal peri-ductular infiltration of mature (small) lymphocytes and neutrophils and there was evidence of a mild to moderate multifocal portal fibrosis. There was also a venous thrombus in one section of the liver. All these observations are illustrated in Figure 9a,b.

**Figure 9.** *Cont.*

**Figure 9.** Microscopic view of the liver (**a**,**b**) (scale bar: 100 μm) and kidney (**c**,**d**) (scale bar: 50 μm) in IP-AF(D)NMs, IP-(D)NMs and IV-drug only treatment groups, stained with H and E. (**a**) Cystic distension of bile ductule containing bile (dark brown thick arrows), foci of hepatocellular necrosis (dark blue arrow), mild bile ductule proliferation (red arrows), and mild lymphocytic infiltration into portal areas (green arrow); (**b**) multifocal to coalescing hepatocellular coagulative necrosis with (red thin arrows) or without associated haemorrhage (red thick arrows) which infiltrated with heterophils (yellow arrow); (**c**) severe nephrosis as evidenced by ragged intraluminal clumps of cell debris (yellow rings), inflamed cells (thick short yellow arrows), thin long yellow arrows denote clusters of heterophils HE; and (**d**) basophilic (calcified), granular intra-tubular cellular detritus (yellow arrow), multifocal (segmental) karyolysis in some tubular epithelial cells, and multifocal mild distension of proximal convoluted tubular lumens (lined by slightly attenuated epithelium), HE (Hematoxylin and Eosin staining).

#### 2.4.8. Renal Histopathology

There was widespread but only mild/moderate cortical and medullary intra-tubular cast formation evident as basophilic, granular, and variably vacuolated (derived from degenerate tubular epithelial cells) cellular detritus that filled some tubules. The outer cortex presented with multifocal (patchy) severe nephrosis-lytic necrosis of proximal convoluted tubular epithelial cells, as evidenced by clumps of ragged, mineralized cell debris (due to dystrophic calcification) and occasional karyolitic nuclei in the affected tubular epithelium. In the deeper cortex, there was widespread desquamation/shedding of proximal convoluted tubules (PCT) epithelial cells into tubular lumens, with slight condensation of nuclear chromatin in the affected epithelial cells. There was mild distension of occasional PCT lumens, which were lined by mildly attenuated epithelial cells. All these observations are illustrated in Figure 9c,d.

#### 2.4.9. Immunohistochemistry (IHC)

MUC16/CA125 IHC revealed positive labeling of 1–10% of the area within the specified regions of interest (ROI). Generally, throughout these tumors, variably-sized clusters of neoplastic cells were labeled with the MUC16/CA125 antibody. Labeling was both cellular (cytoplasmic and cell membrane) and extracellular (around shrunken apoptotic-like cells and in tubular lumina lined by irregularly branching papillae of neoplastic epithelial cells). There was occasional distinct membranous labeling of epithelial cells lining the papillary projections in some sections. However, most labeling appeared extracellular in most tumor sections, and there was more cytoplasmic labeling of neoplastic cells compared to membranous labeling (Figure 10a–c).

**Figure 10.** MUC16/CA125-positive labeling of EOC cells in mouse tissues and tumor foci in IP-inoculated nude mice (MUC 16/CA 125, scale bar: 50 μm): (**a**) Negative tissue control (normal mouse liver). MUC16/CA125 immunolabeling, DAB chromogen, hematoxylin counterstain; (**b**) IP-ovarian carcinoma infiltrating skeletal myofibrils (stars) of the body wall. EOC cell-associated (yellow arrow) and extracellular (red arrows) MUC16-positive labeling. MUC16/CA125 immunolabeling, DAB chromogen, hematoxylin counterstain; (**c**) IP-inoculated ovarian carcinoma infiltrating the pancreas (stars) with MUC16-positive labeling of cytoplasmic membranes (arrows) and extracellular (\* asterisks) labeling of neoplastic cells. MUC16/CA125 immunolabeling, DAB chromogen, hematoxylin counterstain.

In addition, the MTX IP delivery by the polymeric ISFI was intracellularly taken up by OVCAR 5 tumor cells via endocytosis. The intracellular-MTX was released by a lysosomal route and distributed to the cytosol where it attached to its dihydrofolate reductase target enzyme and generated a robust S-phase portion. This modification in ovar5 cell cycle provided the anti-proliferative effect in vivo, demonstrated by the regression in tumor volume.

#### 2.4.10. MUC16/CA125 IHC Analysis on Formalin-Fixed, Paraffin-Embedded (FFPE) Epithelial Ovarian Cancer (EOC) Tissue Sections

The percentage of MUC16-positive labeling per square centimeter of each selected EOC tissue section (one per mouse) was determined with the aid of the phase separation function for both pre and post-treatment mice. Positive labeling was identified as being brown in color and was observed in both cellular (cytoplasmic and cell membrane) and extracellular locations (in tubular lumina) throughout the neoplastic foci (Figure 11A). MUC16 expression in pre and post-treatment groups using both drugs (methotrexate and cisplatin) was up-regulated and displayed an increasing trend in the placebo (4.31–5.11%), AF(D)NMs (6.13–6.36%) and IV-drug groups (7.67–8.14%) (*p* < 0.05), whilst in the NMs (D) group, MUC16 expression was down-regulated (2.67–3.3%) (Figure 11).

AF(D)NMs reduced CA125 antigens, thereby effectively decreasing their expression. This was validated in quantification of plasma and ascitic fluid MUC16/CA125 antigen levels using the Cancer Antigen CA125 Human ELISA Kit (Figure 8). MUC16/CA125 IHC analysis on FFPE EOC tissue sections (as shown Figure 11B) proved not be a satisfactory tool in differentiation of MUC16 expression levels from pre and post-treatment groups, due to the inaccuracy, ambiguity, and superficial nature of this technique.

**Figure 11.** ROI (regions of interest) measurement of IHC images on the stained slides for each treatment group; the MUC16 density in each image was calculated as percentage of MUC16-positive labeling per square centimeter of each selected EOC tissue section (one per mouse). (D) refers to methotrexate (**A**) and cisplatin (**B**). \* *p* < 0.05 by *t*-test.

The survival rate in response to various chemotherapeutic protocols was a significant index for comparing antitumor efficacy between the groups. Nude mice survival rates were significantly different (*p* < 0.05) between the experimental treatment and control/placebo and comparison (IV chemotherapy) groups, as shown by means of Kaplan-Meier analysis. Mouse survival was significantly improved (100% over 35 days) in the AF(D)NMs test group compared to the placebo group, as well as all the other groups (Figure 12). The data also indicated that the mice in the AF(D)NMs test group exhibited the greatest overall reduction in tumor size and had the longest survival times.

**Figure 12.** Kaplan-Meier mouse survival curves, showing the three chemo-treatment groups (*n* = 10) including nude mice injected with IP-AF(D)NM's, IP-(D)NM's and IV chemotherapeutic drugs and an IP-control placebo implant treatment group. The survival rate of mice in the IP-AF(D)NM's test group was significantly improved compared with the other groups. \* *p* < 0.05.

#### 2.4.11. Mechanism of Intraperitoneal ISFI Delivery for Human Ovarian Carcinoma Targeting

The ISFI was fabricated by encapsulating a nanocomposite comprising of anti-MUC16 (antibody) functionalized methotrexate (MTX)-loaded PNIPAAm-b-PASP nanomicelles (AF(MTX)NM's) within a C–P–N hydrogel. Given that the peritoneal cavity is the principal site of disease in ovarian cancer, the ISFI was injected and released nanocomposites into the peritoneal cavity. Following the release of nanocomposites from the hydrogel, the nanomicelles (formulated to circulate for prolonged periods in the peritoneal fluid) targeted specific mucin antigens significantly over-expressed on ovarian cancer cells circulating in the peritoneal fluid (when patients are usually diagnosed) and cancer cells forming nodules at distant sites in the peritoneal cavity. This targeting system reduced the tumor load responsible for adhesion at the sites of secondary metastasis (peritoneal and abdominal surfaces). The anti-MUC16 antibody conjugated nanomicelles has great potential in improvement of tumor selectivity, eliminating/reducing the tumor load, while improving the recovery and long term survival rate of most patients suffering from ovarian cancer.

#### **3. Discussion**

Ovarian cancer is an insidious female disease that is asymptomatic during its early stages and is usually diagnosed at an advanced, often untreatable stage when the disease has spread beyond the ovaries throughout the abdominal cavity and even further afield [1–3]. Thus, deaths due to ovarian cancer could be significantly lowered by developing new, ultrasensitive, yet reliable methods for early diagnosis, and by developing improved treatment protocols. Mucins are among the most promising molecular biomarkers for EOC and have proven invaluable in the diagnosis and monitoring of treatment regimes in various types of ovarian cancer. Increased mucin expression, otherwise known as "mucin switching" that occurs during the neoplastic transformation of ovarian surface epithelium to ovarian carcinoma is important in the progression of this disease [26–28]. MUC16 is an important clinical biomarker of ovarian cancer and it is a target for various immuno-chemotherapies currently under investigation [29–31].

Blood, serum, plasma, and ascitic fluid levels of MUC16 in mice and humans have been well-researched and are now commonly used as diagnostic markers when evaluating the effect of various chemo-treatments and for establishing the advent of relapse in EOC [32,33]. In this study, quantification of MUC16/CA125 concentration was conducted utilizing the highly sensitive Cancer Antigen 125 (CA125) Human ELISA Kit with a date-to-date coefficient of sensitivity >5%. Progressively increasing MUC16/CA125 values are correlated with ovarian malignancy, whilst steady MUC/CA125 values, even when raised, are correlated with some benign conditions. Significantly, in the pre- vs. post-treatment groups of mice with induced EOC, the MUC16 antigen levels in the blood were 2.368 U/mL in the placebo group, compared to 1.9460 U/mL in the (AF(D)NMs) post-treatment group (Figure 11). Our findings also showed that plasma MUC16 antigen levels were consistent with tumor growth, ascitic fluid development and mouse survival with distinct and combination chemo-treatments, indicative of MUC16 as a valuable biomarker to investigate the effect of chemo-therapeutic interventions on EOC in Athymic Swiss mice.

Histopathology was subsequently performed on all IP nodules and in all cases, the existence of anaplastic ovarian carcinoma(s) was confirmed histologically (Figure 10). Cancer linked inflammation response includes white blood cells infiltration, notably phagocytic macrophage cells, the occurrence of polypeptide representatives of inflammation (as tumor necrosis feature, interleukin-1 and chemotatic chemokines) and the presents of tissue alteration and angiogenesis [34]. This flow of episodes presents an uncontrolled pattern that can promote cell laceration and necrosis. Our findings revealed foci of coagulative necrosis but necrotic foci were most often associated with cholestasis and bile infarcts thought to be due to obstruction to bile outflow by tumor nodules. In prior studies, additional toxic side effects that were observed in association with the use of methotrexate/cisplatin included chronic interstitial, obstructive pulmonary disease, severe renal outer cortex nephrosis [35–37]. Immunohistochemical analysis of epithelial ovarian tumors in the present study confirmed an apparent

over-expression MUC16 in all induced tumors. Significantly, cellular labeling of MUC16 was detected both in cytoplasmic granular and/or (to a lesser extent) cytoplasmic membrane labeling of neoplastic cells. Our results showed that MUC16 is up-regulated in most of the tumor tissue samples analyzed (Figure 11). As has been shown in previous studies, we also detected a down-regulation of MUC16 expression in ovarian tumors post-treatment as compared to the pretreatment stage. At euthanasia, MUC16 antigen expression in the ovarian carcinoma tissues significantly decreased post-treatment with the AF(D)NMs and only a few small areas of cell debris stained positive in these sections. MUC16/CA125 IHC analysis on FFPE EOC tissue sections (as shown Figure 11B) proved not be an invaluable tool in differentiation of MUC16 expression levels from pre and post-treatment groups, due to the inaccuracy, ambiguity, and superficial nature of this technique. The improved survival rate associated with the AF(D)NMs treatment was probably at least partly due to the profound decrease in ascitic fluid formation in this group (*p* < 0.01). The reduction in ascitic fluid probably occurred due to the specific and prolonged retention of AF(D)NMs hydrogels intra-peritoneally close to the tumor burden; intra-peritoneal tumor cells had close and prolonged exposure to the chemotherapeutic drugs, no doubt resulting in increased drug efficacy in situ. The novel aspect of the study was the design and implementation of an antibody-functionalized nanomicelle hydrogel composite to specifically target MUC16 antigens known to be over-expressed on ovarian cancer cells. Following the release of antibody-bound nanomicelle from the hydrogel, said nanomicelle targeted specific MUC16 antigens over-expressed on ovarian carcinoma cells within the ascitic fluid and within neoplastic nodules at distant sites within the peritoneal cavity. This MUC16 targeting system reduced the mass of neoplastic cells capable of adhering to parietal and visceral peritoneal surfaces at sites distant to the original tumor implant. Chemo-treatment with AF(D)NMs significantly inhibited growth of and transcoelomic metastasis of EOC (via reducing MUC16 antigen expression on tumor xenografts) and reduced the production of ascitic fluid, thereby possessing potential for increasing longevity in patients with ovarian carcinoma.

#### **4. Materials and Methods**

#### *4.1. Materials*

Natural and synthetic biopolymers such as Chitosan (MW = 600.000) and Poly(*N*-vinylpyrrolidone) (MW = 40.000) were purchased from Sigma-Aldrich (Steinhelm, Germany). *N*-isopropylacrylamide (MW = 113), poly(*N*-vinylpyrrolidone) (MW = 40.000), *N*,*N*-methylenebisacrylamide, glutaraldehyde (GA) (25%), 2,2 -azoisobutyronitrile, ammonium persulfate, α-amino Aspartic acid, ethanethiol hydrochloride (AET-HCl), 98% analytical grade triphosgene, methotrexate (MTX), 3.5 mg/mL penicillin, 0.1 g/L streptomycin, RPMI 1640 medium, 10% FBS and 0.25% trypsine, and 0.03% EDTA solution were procured from Sigma Aldrich (St. Louis, MO, USA). 98% DMSO, 99.8% anhydrous *N*,*N* -dimethylformamide (DMF), 99.0 anhydrous tetrahydrofuran (THF), 98.0% diethyl ether, 75% analytical petroleum ether bp (30–60 ◦C), 1 M hydrochloric acid, potassium chloride (KI), disodium hydrogen phosphate (Na2HPO4), and potassium dihydrogen phosphate (KH2PO4) were from Merck Chemicals (Pty) Ltd. (Darmstadt, Germany). NIH:OVCAR-5 cancer cells were procured from director of Fox Chase Cancer Institute, Philadelphia, PA, USA (Dr. Tom C. Hamilton). RayBio® Human CA-125 (MUC16) Elisa Kit for serum, plasma, cell culture media, and urine 96-wells filter plates were procured from Biocom Biotech in (Centurion, Pretoria, South Africa). Anti-MUC16 antibody [OC125] ab693 was procured from Abcam Inc. (Cambridge, MA, USA). Purified deionized water was prepared by a Milli-Q System (Millipore Co., Billerica, MA, USA). Culture plates were purchased from Corning Inc. (Corning, NY, USA). Ovarian cancer cells were cultured in a RS Biotechnological Galaxy (Irvine, UK) incubator kept at 37 ◦C with 5% carbon dioxide in an entirely humidified environment. Cell experimentations were conducted in the logarithmic stage of development.

#### *4.2. Anti-MUC16 Functionalized MTX-Loaded NanoComposite Preparation*

#### 4.2.1. Synthesis of PNIPAAm-b-PAS for NanoComposite Formulation

Firstly, a copolymer comprising PNIPAAm-b-PASP was synthesized by solvent evaporation as described previously [38,39]. Thereafter the PNIPAAm-b-PASP copolymer (0.125 g, 0.0079 mmol), anti-MUC16 antibody (0.2 mL, 0.022 mmol), NHS (0.60 mg, 0.0522 mmol), and DCC (10.8 mg, 0.0522 mmol) were dissolved in 10 mL DMF. The solution was mixed under inert conditions at room temperature in the dark for 14 h before dilution with 25 mL deionized water followed by centrifugation to extract DCU. The aqueous supernatant was further extracted by membrane dialysis against deionized water for 24 h with subsequent lyophilization.

#### 4.2.2. Preparation of the MTX-Loaded NanoComposite

MTX-loaded nanomicelles were prepared by solvent evaporation as reported previously [31]. Briefly, 7 mg MTX and 25 mg of PNIPAAm-b-PASP were dissolved in 5 mL DMF and 20 mL distilled water was added drop wise to the solution under homogenization at 360 rpm for 5 min at room temperature. This was followed by solvent removal using a rotary vacuum evaporator to obtain a nanomicelle solution that was then filtrated through a 0.2 μm filter membrane to remove any residual MTX with subsequent lyophilization. The anti-MUC16 antibody functionalized MTX-loaded nanomicelles were prepared with modifications by introducing the anti-MUC16 functionalized block copolymer [40,41].

#### 4.2.3. Evaluation of the Molecular Structural Integrity of the Functionalized Nanomicelles

Fourier Transform Infrared (FTIR) spectroscopy (Perkin Elmer Life spectroscopy 100, Lantrisant, Wales, UK) was utilized to differentiate the chemical structure of the block copolymer NanoComposite following nanomicelle coating with the anti-MUC16 antibody, employing a oscillating vibrational component with a diamond interior gauge. Evaluation of Pure MTX, blank nanomicelles, MTX-loaded nanomicelles and the MTX-loaded NanoComposite was conducted at an interval of 650 and 4000 cm−1 wavenumber with a 4 cm−1 resolution value and smoother 64 scans per spectrum.

#### 4.2.4. Differential Scanning Calorimetry for Elucidation of the Thermal Events of the Methotrexate-Loaded Nanomicelles

Differential Scanning Calorimetry (DSC) studies were conducted with a Mettler Toledo advanced DSC1-STARe equipment from Schwerzenback in Switzerland. The Mettler-STARe software program (version 9, Schwerzenback, Switzerland) was utilized for DSC results acquisition and interpretation. Blank specimen pans were used as reference points and experimental scans were conducted by heating samples at intervals of −10 to 25 degrees for fifteen minutes with a continual isotherm. The lyophilized powder specimens (10 mg) were placed into DSC aluminium-pans and compress-sealed. Relative DSC scans were conducted on blank micelles, MTX-loaded micelles and pure MTX at a heating rate of 10 ◦C/min from −10 to 325 ◦C in a stable stream of nitrogen gas. Fresh sample analysis was conducted in triplicate for method development and validation purposes. DSC thermogram patterns were then analyzed for variations in thermal periods. The phase changes of the drug-free micelle and MTX were correlated with the phase changes of the MTX-loaded PNIPAAm-block-PAsp nanocomposites.

#### *4.3. Chitosan-PVP-PNIPAAm (C–P–N) Hydrogel Synthesis*

Interpenetrating polymer networks (IPN's) were fabricated by free-radical polymeri-zation mechanism of *N*-isopropylacrylamide distinctive monomers in excipient chitosan and water-soluble poly(*N*-vinylpyrrolidone) (PVP). PVP, *N*-isopropylacrylamide NIPAAm, crosslinkers-*N*,*N* -methylenebis(acrylamide)-MBAAm, and glutaraldehyde dissolved in aqueous chitosan-CHT (0.02 L, 2%, 1.6% CH2COOH-acetic acid), then stirred. CHT, PVP, NIPAAm employed were in 2:2:45 mass percentage. The percentages (*w*/*w* %) between glutaraldehyde-OHC(CH2)3HO

and CHT were 1%, 2%, 4%. 4% weight percentage existed between MBAAm with NIPAAm *N*,*N*,*N* ,*N* -Tetramethylethylenediamine (TEMED)/ammonium persulfate (APS) solution was utilized as initiator in the cross-linking process of NIPAAm. Thereafter, APS-TEMED solution (24 μL, 4%) was subsequently added to this reaction mixture. A cross-linked Chitosan**-**PVP-PNIPAAm (C–P–N) hydrogel formed within 1 h of polymerization reaction at 25 ◦C. After 24 h, the IPN hydrogels were cleaned with deionized H2O, then air/vacuum dried. The hydrogel samples prepared were called F1 = C–P–N/1, F2 = C–P–N/2, and F3 = C–P–N/3 IPNs.

#### 4.3.1. Nuclear Magnetic Resonance (NMR) Spectroscopic Analysis

I 1H-NMR measurements were applied to the C–P–N components for the confirmation of the copolymer structure and composition. For 1H-NMR measurement, 5 mg sample vacuum dried at 50 ◦C for 48 h was added into a 5 mm NMR test tube, and further vacuum dried at 50 ◦C for 48 h, to which 500 μL D2O solvent was introduced, and lastly the test tube was vortexed to dissolve the polymer in solution. The NMR spectrum was generated using a Bruker DRX400 spectrometer (Bruker, Germany).

#### 4.3.2. Determination of Polymeric Structural Variations

Molecular structural changes in the polymer backbone may alter the inherent chain stability and therefore affect the physicochemical and physicomechanical properties of the selected polymer type for the intended purpose. The molecular structure of native polymers (CHT, PVP, PNIPAAm) the non-cross-linked, and cross-linked C–P–N hydrogel, blank micelles, and drug-loaded micelles were analyzed using FTIR spectroscopy to elucidate any variations in vibrational frequencies and subsequent polymeric structure resulting from drug-co-polymer interaction during nanostructure and hydrogel formation. Samples were analyzed in triplicate at high resolution with scans ranging from 4000 to 400 cm−1 on a PerkinElmer Spectrum 100 Series FTIR spectrometer coupled with Spectrum FTIR research grade software (Perkin Elmer Life and Analytical Sciences Inc., Hopkinton, MA, USA).

#### *4.4. Preparation of Bio-Responsive IPN Nanomicelle/Hydrogel Composite Based Implant (ISFI)*

The ISFI was fabricated by encapsulating antibody functionalized nanocomposite comprising anti-MUC16 (antibody) functionalized methotrexate (MTX)-loaded PNIPAAm-b-PASP nanomicelles within the synthesized bio-responsive C–P–N hydrogel followed by freeze drying. Anti-MUC16 functionalized nanocomposites were encapsulated into the C–P–N at a proportion of 1:5 (NM's: C–P–N hydrogel). The anti-MUC16 functionalized nanocomposites in solution were introduced drop-wise into the C–P–N gel, and syntheses were permitted to stir until a uniform mixture was reached with a mechanical vortexing system at 37 ◦C.

#### 4.4.1. Determination of the Gelation Temperature of the Polymeric ISFI Utilizing Oscillatory Rheology

Dynamic rheology is one of the most extensive methods to study rheological properties of polymer hydrogels, and is also the most direct and reliable way for the determination of sol–gel transitions. Viscoelastic properties were measured with a Modular Advanced Rheometer system (ThermoHaake MARS Rheometer, Thermo Fischer Scientific, Karlsuhe, Germany) to indicate the storage modulus (G ), the loss modulus (G"), and tan δ of an aqueous solution of the sol, using cone plate geometry, where the G and G" were recorded under constant deformation. The study of the flow properties was considered extremely significant to this study as it is central to the mechanisms by which the ISFI functions. At room temperature, the implant remains in the liquid state to allow delivery via an 18 G needle and at body temperature (37 ◦C) the implant forms a solid-like structure. To characterize and analyze the flow behavior and determine the lowest critical solution temperature (LCST) or gelation temperature of the ISFI formulations, rheology studies were conducted using a Haake Modular Advanced Rheometer System (ThermoFisher Scientific, Karlsruhe, Germany). As the polymeric material acts as a visco-elastic solid, some background on viscoelastic solids is provided. To determine the lower critical solution temperature (LCST) and hence the gelation temperature of the ISFI formulation, the temperature of

samples was ramped from 20–50 ◦C at a rate of 0.25 ◦C/min while applying the predetermined stress of 13.4 Pa, at the frequency of 10 Hz. The gelation temperature was determined as the temperature at which the cross-over of G and G" occurred, i.e., the point at which the formulation was no longer acting as a liquid (G"), but as a solid (G ). In all cases, a solvent trap was used to prevent sample evaporation.

#### 4.4.2. Morphological Characterization of the ISFI

Peripheral morphology, the surface area, and cross sectional area of the ISFI post-lyophilization were observed by using scanning electron microscopy—SEM (JEOL Ltd., JSM-Japanese-Electronic-Optical-Laboratories, Tokyo, Japan). Microscopic observations of the exterior-surface of the nanocomposite configuration of the C–P–N hydrogel was conducted, by first freeze drying the hydrogel at 25 mTorr pressure (Virtis™-Gardiner, New York, NY, USA). The specimen was attached to dual-sided adhesive tape adhered to a metallic specimen stand and sputter-covered with a film of gilded gold. Each specimen was observed under different increasing magnifications at 20 keV-an accelerating-voltage.

#### *4.5. In Vivo Studies in the Swiss Athymic-Nude Mice Model*

#### 4.5.1. Mouse Model

Four- to six-week-old female Swiss Athymic nude mice (purchased from Charles River Laboratories International Inc., Ecully, France) were housed in a specific pathogen free (SPF) facility at 25 ◦C with 60–70% relative humidity (RH) with a 12-h light/dark rotation and then given a week of acclimatization before the experiments commenced. Mice were fed ad libitum diet and water. All procedures were performed under sterile conditions in a laminar flow hood. The animals were monitored daily for general health status. Animal ethics for in vivo studies was approved by University of the Witwatersrand, Animal Ethics Screening Committee (AESC Number: 2012/46/05; Dated 5 December 2012) and the study protocol adhered to these guidelines and those of the South African Council on Animal Care.

#### 4.5.2. In Vitro Cell Culture

The NIH:OVCAR-5 ovarian cancer (OC) cell line is a developed human ovarian epithelial carcinoma cell line that transmits the MUC16 autoantigen [26]. These NIH:OVCAR-5 cells were procured from the director of the Fox Chase Cancer Facility, USA-Jenkintown-CA (T. C. Hamilton). The cells were developed to 80% confluency and attached in plastic culture containers in RPMI growth medium augmented with 10% heat-inactivated FBS, 2 mM glutamine, 50 units/mL penicillin/streptomycin, and 50 units/mL insulin. OC cells were cultured in a RS Biotechnological Galaxy (Irvine, UK) incubator kept at 37 ◦C with 5% carbon dioxide in an entirely humidified environment. OC cells were weekly sub-cultivated, collected by dissociation of adherent cells with trypsin, twofold cleaned in 10% PBS, dead cells tainted with sterile Trypan Blue and enumerated. Cell culture experiments were performed after 1–2 sub-cultivation rotations. An aqueous 0.4% *w*/*v* trypan blue in sterile 10%, 0.1 M, pH 7.4 PBS was formulated for cell enumeration. A sterile hemocytometer grid was the tool employed for enumeration and a 20 μL cell mass and 60 μL trypan blue volume were collectively mixed. The ratio of live cells was computed as a determinant of the amount of cells enumerated per total number of grid quadrants counted as depicted in % live cell Equation (1). This was 1:3 proportion which generated a 4 dilution-factor.

$$\% \text{ Live cells} = \frac{\text{No of cells counted}}{\text{No of grid quadrants counted}} \times DG \times 10^4 \text{ cells/mL} \tag{1}$$

where *DG* is the dilution-factor utilized and 104 is equation constant. Only specimens that demonstrated viability of more than 95% were used in following analysis.

#### 4.5.3. Induction of Human Ovarian Carcinoma in Swiss Athymic Nude Mice—Pre-Treatment Phase

The ovarian carcinoma induction stage of this study was referred to as the pre-treatment phase. All experiments were conducted inside an Esco Frontier™ DuoFume Hood in the specific pathogen free (SPF) facility. In the present, the model was generated by injecting female Athymic 4- to 6-week-old nude mice IP with 2 × 108 cells/mL of NIH:OVCAR-5 cancer cell suspensions in 0.2 μL RPMI. Inoculations were performed using a 26 G gauge needle and a 1 mL syringe [26]. Human ovarian carcinomas (NIH:OVCAR-5) were induced within 10 days of inoculation. As NIH:OVCAR-5 mainly proliferated within the peritoneal cavity and over-expressed mucin 16 (MUC16/CA125) as compared to normal tissue, it was concluded that the IP growth mimicked human clinical disease. All tumor-bearing mice were further monitored until the tumors developed to their target size of 80–100 mm3, which was regarded as the baseline-for initiation of treatment [27].

#### 4.5.4. IP-Inoculated Mice