Design and Characterization of Endostatin-Loaded ...downloads.hindawi.com/journals/jnm/2017/2539065.pdfF11 0.55 0.3 2.5 56.28 +9.08 0.055 1.000 F12 1.00 0.2 0.5 329.10 −12.30 0.052

Research ArticleDesign and Characterization of Endostatin-LoadedNanoparticles for In Vitro Antiangiogenesis in SquamousCell Carcinoma

Samson A. Adeyemi, Yahya E. Choonara, Pradeep Kumar, Lisa C. du Toit, and Viness Pillay

Wits Advanced Drug Delivery Platform Research Unit, Department of Pharmacy and Pharmacology, School of Therapeutic Science,Faculty of Health Sciences, University of the Witwatersrand, Johannesburg, 7 York Road, Parktown 2193, South Africa

Correspondence should be addressed to Viness Pillay; [email protected]

Received 4 February 2017; Revised 22 March 2017; Accepted 9 April 2017; Published 4 June 2017

Academic Editor: Yan Zou

Copyright © 2017 Samson A. Adeyemi et al. This is an open access article distributed under the Creative Commons AttributionLicense, which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properlycited.

The aim of this study is to effectively enhance antitumor activities of endostatin by preparing polymeric nanocarriers. NMR and FT-IR spectra confirmed the successful grafting of the CHT-g-PEI andCHT-g-PEI-PEG-NH2 conjugates. SEMmicrographs confirmedthe shape of endostatin-loaded nanoparticles to be spherical while both TEM and zeta size results showed nanoparticle’s averagesize to be 100.6 nm having a positively charged surface with zeta potential of 7.95mV. The concentrations of CHT and TPP aswell as the changing pH conditions account for the increased swelling pattern of endostatin-loaded nanoparticles and influencedendostatin release in vitro. PEI increased the overall amine protonationwhile PEG aggravated endostatin encapsulation and release.Nanoparticles swell and release endostatin at acidic tumor pH of 6.8 compared to physiological pH of 7.4. The native CHT-g-PEI-PEG-NH2 conjugate showed high cytocompatibility above 80% cell viability across tested formulations. Endostatin-loadednanoparticles showed a significant reduction in cell viability across tested formulations, with 5.32% cell death at 125𝜇g/mL and13.36% at 250 𝜇g/mL following 24 hours’ incubation period. Interestingly, more than a fourfold (61.68%) increment in cytotoxicitywas observed at nanoparticle concentration of 1000𝜇g/mL. It was concluded that CHT-g-PEI-PEG-NH2 is an effective cargo forendostatin delivery with antiangiogenic effect in squamous cell carcinoma.

1. Introduction

Squamous cell carcinoma (SCC) accounts for different cancertypes emanating from various tissues of the body includ-ing the skin, oesophagus, head and neck, urinary blad-der, prostate, lung, vagina, and cervix resulting in about2,500 deaths annually in the United States [1]. Over theyears, chemoradiation and surgery had been the prominenttreatments options for patients diagnosed with SCC. Untilrecently, with the advent of cutting edge approaches usingnanotechnology, chemotherapeutics employed in the man-agement of SCC were hampered with diverse challengessuch as suboptimal dosage, cytotoxicity to normal cells dueto nontargeted delivery, short circulation time as well asmultiple resistances due to the Reticuloendothelial System(RES) [1].

To date, anticancer drugs have been encapsulated intodiverse cargos such as polymeric micelles, surfaced-modified

particles, liposomes, and nanoparticles for the delivery ofanticancer drugs in cancer nanomedicines [2–4] which havenot been able to overcome these challenges. Meanwhile,nanoparticles, in contrast to other delivery vehicles, have anumber of chemotherapeutic advantages including ease ofinjection, high drug-loading ratio, reduced toxicity to healthycells/tissues and enhanced direct targeting effect in bothprimary and metastatic tumors [2].

Angiogenesis, the formation of new blood vessels, isfundamental to the survival and growth of tumor cells [5–7]. Recent focus has been on the use of natural and syntheticinhibitors of angiogenesis that can prevent or slow down thegrowth of tumor cells by blocking the formation of new bloodvessels as a promising strategy for tumor therapy [5].

Endostatin is a proteolytic C-terminal fragment of colla-gen XVIII with a molecular weight of 20 kDa. Among otherangiogenic inhibitors, endostatin has received the greatestattention for its broad-spectrum and low toxicity [8]. These

HindawiJournal of NanomaterialsVolume 2017, Article ID 2539065, 17 pageshttps://doi.org/10.1155/2017/2539065

https://doi.org/10.1155/2017/2539065

2 Journal of Nanomaterials

Table 1: Box-Behnken experimental design for formulation variables and responses at pH 7.4 and 6.8, respectively.

Formulation

Variables Responses

Surfactant conc.(mg/mL)

TPP conc.(mg/mL)

Polymerconjugate conc.

(mg/mL)

Size(nm)

Surface charge(mV)

Conductivity(mS/cm)

Polydispersityindex

F1 1.00 0.1 1.5 65.07 +16.10 0.054 0.259F2 0.10 0.1 1.5 84.22 +19.80 0.079 0.367F3 0.10 0.2 2.5 173.9 +10.40 0.056 0.689F4 1.00 0.3 1.5 97.38 −1.87 0.048 0.302F5 0.55 0.2 1.5 300.10 +11.92 0.053 0.529F6 0.55 0.2 1.5 295.10 +9.05 0.053 0.715F7 0.55 0.1 2.5 163.60 +11.00 0.038 0.268F8 0.55 0.3 0.5 169.10 +8.95 0.062 0.731F9 0.55 0.1 0.5 125.20 −6.11 0.067 0.403F10 1.00 0.2 2.5 125.50 +0.76 0.049 1.000F11 0.55 0.3 2.5 56.28 +9.08 0.055 1.000F12 1.00 0.2 0.5 329.10 −12.30 0.052 0.788F13 0.10 0.2 0.5 127.21 −8.94 0.057 1.000F14 0.10 0.3 1.5 363.10 −7.63 0.062 0.408F15 0.55 0.2 1.5 310.20 +13.10 0.048 0.550

advantages speed up the investigation process of endostatininto the clinical trial [9]. Endostatin has been shown to inhibitendothelial cell proliferation, migration, and formation ofnew blood vessels [7, 10]. However, as any other protein,endostatin has many clinical challenges in its applicationsuch as high dosage to maintain its efficacy, high price,short half-life, and instability [11]. As such, incorporation intobiodegradable polymers as delivery cargos could amelioratethese limitations.

Chitosan (CHT), as opposed to other naturally occur-ring polymers and CHT-based nanoparticles, has recentlyattracted much consideration in both pharmaceutical andbiomedical applications due to its exceptional biologi-cal properties including biocompatibility, biodegradability,and nontoxicity [12] but with low transfection efficiency.The high buffering potential and transfection efficiency ofpolyethylenimine (PEI) have been explored for the deliveryof DNA and other anticancer therapeutics in the man-agement of cancer diseases [13–15]. Meanwhile, covalentattachment of hydrophilic polyethylene glycol (PEG) ontothe surfaces of nanoparticles prolongs the circulation half-life in vivo of encapsulated chemotherapeutics, shields thesurface of nanoparticles from uptake by the RES, andreduces carrier’s cytotoxicity with improved colloidal stability[16].

In this study, we employed low molecular weight CHTgrafted onto PEI as a cationic carrier for improved deliveryof endostatin. Furthermore, the surfaces of these gratedpolymerswere coatedwithPEG to improve endostatin encap-sulation and prolong release in vitro. Surface PEGylationof the grafted polymer also enhances its stability in theextracellular matrix of oesophageal squamous cell carcinoma(OSCC) cells. Formulation parameters, optimization andcytocompatibility of the nanosystem were evaluated and the

antiangiogenic effect of endostatin-loaded nanoparticles wasassessed in an OSCC cell line (KYSE-30).

2. Materials and Methods

2.1. Materials. Low molecular weight chitosan (DA = 75–85%, MW = 50–190 kDa), branched polyethylenimine (PEI)(Molecular weight Mw = 25KDa), Bovine Serum Albumin(BSA) (MW=66 kDa), human recombinant endostatin (MW= 22 kDa), 1-ethyl-3-(3-dimethylaminopropyl) carbodiimidehydrochloride (EDC), N-hydroxysuccinimide (NHS), so-dium tripolyphosphate (TPP) (MW = 367.86 g/mol), poly(vinyl alcohol) (PVA) (MW = 85,000 g/mol), tricholroa-cetic acid (TCA), and acetonitrile (ACN) were purchasedfrom Sigma-Aldrich Co., Ltd. (St. Louis, MO, USA). Func-tionalized poly(ethylene) glycol (NH2-PEG-COOH, Mw =2100 g/mol) was from NANOCS (New York, NY, USA). Celllines KYSE-30, RPMI, HAM’s F12, Fetal Bovine Serum (FBS),and pentamycin/streptomycin were from Life Bioscience(Oakleigh, VIC, Australia). All other solvents and reagentswere of analytical grade unless stated otherwise.

2.2. Preparation of [CHI-g-PEI-PEG-NH2]-Endostatin-Loaded Nanoparticles

2.2.1. Synthesis of the CHI-g-PEI Conjugate. A Box-Behnkenexperimental design was employed to generate fifteennanoformulations as presented in Table 1. Amodifiedmethoddescribed by Gao et al. [17] was employed to synthesizethe CHT-g-PEI conjugate. Briefly, 0.5, 1.5, and 2.5mg/mLchitosan (CHT) solutions were prepared in 0.5% aceticacid and left overnight (Table 1). 1,1-Carbonyldiimidazole(CDI) was added to the CHT solutions and stirred for 1hour at room temperature to activate the amine group in

Journal of Nanomaterials 3

the CHT solution. Subsequently, 0.25% v/v of branchedpolyethylene imine (PEI) was gently added using a needleand syringe at a molar ratio of CHT amine : PEI concen-tration of 2 : 1. The reaction mixture was left to polymer-ize for 24 hours and dialyzed using a dialysis membrane(MW = 12,000 kDa) over double deionized water (DDW)for 24 hours. The final CHT-g-PEI powder was then col-lected by lyophilization over 24 hours (details of equipmentused).

2.2.2. Synthesis of CHI-g-PEI-PEG-NH2 Conjugate. Copoly-mer synthesis procedure described by Jiang et al. [18]with modifications was employed in the synthesis of theamino terminal CHT-g-PEI-PEG-NH2 copolymer conju-gate following an amide formation reaction between theactivated carboxyl groups of NH2-PEG-COOH and theamine groups of CHT-g-PEI as previously described. Briefly,the carboxyl group of the bifunctional PEG (NH2-PEG-COOH) was activated using NHS/EDC chemistry for 15minutes. Furthermore, the CHT-g-PEI conjugate previouslysynthesized was gently added to the activated NH2-PEG-COOH solution using a needle and syringe under stirring.The reaction was then allowed to polymerize at 4∘C and25∘C for 12 hours respectively. The synthesized CHT-g-PEI-PEG-NH2 conjugate was then dialyzed using a dialysismembrane (MW = 12,000 kDa) against DDW for 48 hoursand the dried power was collected by lyophilization over 24hours.

2.2.3. Bovine Serum Albumin/Endostatin Entrapment andNanoparticle Synthesis. An ionotropic gelation techniquewas employed for the synthesis of bovine serum albuminor endostatin-loaded nanoparticles comprising the fifteenformulations and the optimized nanosystem, respectively.The fifteen formulations were prepared according to theformulation variables presented in Table 1. For the experi-mental design formulations, bovine serum albumin (BSA)was employed as a model drug for protein therapeutics foridentification of the optimumnanoparticle system intowhichendostatin would ultimately be loaded. Briefly, 1mg/mL BSAsolution was mixed with the solution mixture of CHT-g-PEI-PEG-NH2 conjugate under mild stirring for 10 minutes.Varying concentrations of TPP as a cross linker were addeddrop-wise to the BSA-loaded mixture using a needle andsyringe. The formation of an opaque and turbid solutionconfirmed the formation of the BSA-loaded nanoparticles.Varying concentrations of PVA were also added as surfactantduring the nanoparticle synthesis as shown in Table 1. Thesynthesized nanoparticles were then allowed to undergo gela-tion for 1 hour under mild stirring and the resultant gel wascentrifuged at 5000 rpm for 1 hour.The clear supernatant wasdiscarded and the pellet was resuspended inDDWand refrig-erated at −80∘C. The dried powder BSA/endostatin-loadednanoparticles were then collected following lyophilizationover 24 hours.

2.3. Analysis of Chemical and Functional Transformationsvia Fourier Transform Infrared Spectroscopy. The vibra-tional transitions in the chemical structures of CHT-g-PEI,

CHT-g-PEI-PEG-NH2, and endostatin-loaded CHT-g-PEI-PEG-NH2 were evaluated using Fourier transform infrared(FT-IR) spectroscopy (PerkinElmer Inc., Waltham, Mas-sachusetts, USA). Samples were placed on a single bouncediamond crystal and processed by a universal attenuated totalreflectance (ATR) polarization accessory, at a resolution of4 cm−1, with the spectrum ranging from 4000 to 650 cm−1.

2.4. Evaluation of Structural Modifications via Nuclear Mag-netic Resonance Spectroscopy. Nuclear Magnetic Resonance(NMR) spectra were recorded for CHT-g-PEI and CHT-g-PEI-PEG-NH2 conjugates. Proton NMR chemical shifts,expressed in ppm and analyzed in deuterated water (D2O)doped with deuterated acetic acid (CD3COOD) in a ratio of5 : 1 were recorded on a 500MHz Avance III spectrometer(Bruker BioSpin GmbH, Germany) at room temperature.

2.5. Particle Size, Surface Charge, Conductivity, and Struc-tural Morphology Measurement. Malvern Zetasizer Nano ZS(Malvern Instruments, Worcestershire, UK) was employedto evaluate the size, zeta potential, and the conductivity ofthe nanoparticles. Known sample weights were dispersed inDDW and sonicated for 30 seconds. 2mL dispersed sample(2mL) was placed in disposable cuvettes and the dynamicscatter intensity was recorded at 25∘C. The average zeta size,polydispersity index (PDI), zeta potential, and conductivityof the nanoparticles were recorded.

The size of the nanoparticles was also confirmed usingTransmission Electron Microscopy (TEM) (FEI Tecnai T12TEM, 60–120 kV, Hillsboro, OR, USA).The nanoparticle sys-temwas dispersed inDDW, ultrasonicated for 15minutes anda single drop of the nanoparticle suspensionwas dropped on aFormVar� coated 200-mesh copper grip (TAABLaboratoriesEquipment Ltd., Aldermaston, England) and allowed to airdry at 25∘C prior to TEM analyses.

Themorphology of the nanoparticles was examined usingScanning Electron Microscopy (SEM). Powdered sample ofthe nanoparticles was placed onto an aluminium specimenstubcovered with a double-sided carbon adhesive disc andsputter-coated with both palladium and gold for 4 minutes at20 kV. SEM images of endostatin-loaded nanoparticles werethen viewed by a Scanning Electron Microscope (SIGMAVP, ZEISS Electron Microscopy, Carl Zeiss Microscopy Ltd.;Cambridge, UK).

2.6. Determination of the Drug-Loading Capacity and Entrap-ment Efficiency of the Nanoparticles. Known amounts (3mg)of drug-loaded nanoparticles of the 15 formulations and theoptimized nanosystems was dispersed in 5mL DDW andshaken vigorously. The mixture was then centrifuged for 1hour at 5000 rpm. The clear supernatant solution (2mL)was poured into the UV cuvette and the absorbance of thedrug was read at 280 nm for all samples using a NanoPho-tometer. Measurements were performed in triplicate andthecorresponding drug concentrations were computed from astandard calibration curve. The drug-loading capacity (LC)and entrapment efficiency (EE) of the nanoparticles werecalculated using (1) and (2), respectively:


LC =Total amount of drug loaded − Unloaded drug in the supernatant

Measure weight of drug loaded nanoparticles, (1)

EE =Total amount of drug loaded − Unloaded drug in supernatant

Total amount of drug loaded. (2)

2.7. Determination of the Degree of Swelling of the Nanoparti-cles. The swelling behaviour of the synthesized nanoparticleswas determined at both pH 6.8 and 7.4 using the MalvernZetasizer Nano ZS (Malvern Instruments, Worcestershire,UK). Known amounts of nanoparticles were incubated at37∘C in phosphate buffered saline (PBS, pH 6.8 and 7.4) over4 hours.The average size of nanoparticles (nm) was recordedat 0, 1, 2, and 4 hours. Each sample was sonicated for 30seconds (SONICS Vibra Cell�, Newtown, CT, USA), placedin disposable polystyrene cuvettes and the dynamic scatterintensity was recorded at 25∘C for the corresponding pH atthe specific time interval.

2.8. In Vitro Drug Release and Chromatographic Appara-tus/Condition for Reverse Phase HPLC. The drug releaseexperiments were carried out at 37∘C in an orbital shakerincubator (YIHDER, Taiwan) over 16 hours. Equivalentamounts of the BSA-loaded and endostatin-loaded nanopar-ticles were placed in a 15mL centrifuge tube with 5mLPBS buffer at both tumoral pH of 6.8 and physiologicalpH of 7.4. The release buffer medium (2mL aliquot) wasremoved at predetermined time intervals (1, 2, 4, 6, 8, and16 hours), filtered, and replaced with fresh amount of thebuffer medium to maintain sink conditions. UV absorbanceof the filtered samples was thenmeasured at 280 nmusing theAlliance HPLC system, including Waters 2695 Separationsmodule Waters 996 PDA Detector; Millennium 32 chro-matography analysis software, Waters Symmetry 300� C18(4.6mm × 250mm, 5 𝜇m), Waters 1525 binary pump, andWaters 2489UV/visible detector, Superdex� 75Nhr 10/30column. The fixed phase is Symmetry 300TM C18 (4.6mm× 250mm, 5 𝜇m). Mobile phases, solution A is 0.1% TCA-water (1 : 1000), solution B is 0.1% TCA-95% CAN (1 : 1000);flow rate is 1.0mL/min. The temperature of the column orsample plate was set at 30∘C and samples were quantified bythe Lowrymethod.The cumulative BSA or endostatin releasewas measured as a function of time.

2.9. Optimization of the Formulatory Components Using Box-Behnken Experimental Design. Polynomial equations relat-ing the dependent and independent variables using MinitabStatistical Software (MINITAB�, V15, Minitab, USA) wereemployed to calculate the optimized delivery system fromthe formulation process under constrained conditions forthe measured responses. Using the optimization parameters,optimized endostatin-loaded nanoparticles were preparedfollowing the same method employed for the 15 formula-tions. However, 0.5mg/mL human recombinant endostatinwas loaded into the nanoparticles. For generation of theoptimum formulation, the average nanoparticle size as well

as the in vitro cumulative endostatin release from the loadednanoparticles at pH 7.4 was minimized, while the releaseat tumor pH of 6.8 and the overall average surface chargeof the nanoparticles was maximized. Response analysis wasperformed for all the variables tested. Surface and contourplots derived from the responses were used to authenticatethe effects of the independent variables on the responsefactors.

2.10. Nanoparticle Cytocompatibility and Cell ProliferationAssay. In vitro cell cytotoxicity was determined using Ala-mar blue assay. Human oesophageal squamous cell carci-noma cell line (KYSE-30) (Life Bioscience) were seeded incomplete media comprising RPMI and Ham’s F12 (1 : 1), sup-plemented with 10% Fetal Bovine Serum, 2mM glutamine,sodium bicarbonate, and 100 𝜇L penicillin/streptomycin(Sigma-Aldrich; St. Louise, MO, USA). The cells were main-tained in an incubator (RS Biotech Galaxy, Irvine, UK),with a humidified atmosphere of 5% CO2 at 37

∘C. KYSE-30 cells were diluted in a complete medium at a finalconcentration of 5 × 104 cells/mL and seeded (25 𝜇L/well)and incubated for 24 hours prior to cell proliferation eval-uation using Alamar blue assay according to the manu-facturer’s procedures. Native nanoparticles and endostatin-loaded nanoparticles were dissolved in the serum-free culturemedium at varying concentrations of endostatin (125𝜇g/mL,250 𝜇g/mL, 500𝜇g/mL and 1000𝜇g/mL). Attached cells inthe wells were treated with the various nanoformulations intriplicate for 24 hours and cell viability was quantified atmaximum emission/excitation wavelengths of 535 nm and595 nm, respectively, on a microplate reader (FilterMax� F5Multi-Mode Microplate Reader, Molecular Devices, USA).Attached cells treated only with serum-free media, withoutthe nanoformulation, were used as a control. Results arepresented as percentage cell viability (% CV ±mean standarddeviation), with the percentage of viable cells was calculatedusing the following Equation (3).

CV

=Fluorescence reading in treated cells

Fluorescence reading in control (untreated) cells.(3)

3. Results and Discussion

3.1. Polymer Grafting and Nanoparticle Synthesis and Char-acterization. As presented in the reaction schematics inFigure 1, the amine in acetylated low molecular weight CHTwas grafted to the amine of 25KDa branch PEI through theCDI linkage chemistry. The two imidazolyl groups of CDI


OH

OH

HONH

OH

HO HO

4

5 2

3 1

6

Chitosan (CHT)

CHT-g-PEI

EDC/NHS

CDI

Branched PEI

y y

y

n

n

n

n

O O

O

O

O

O

O

N N

HN

N

N

NNN NH

HN

N

NN

N N

O

OO

O

OHOOC

NH

NNH

H

NH2

NH2

NH2

NH2

NH2

NH

NH

NH

NH

HN

N N N

N

NH

HN

HN

HN

NH2

COOH-PEG-NH2

NH2

NH2

NH2

NH2

NH2

CHT-g-PEI-PEG-NH2

H2N

H2N

H2N

H2N

H2N

+

Figure 1: Proposed reaction scheme for the preparation of grafted polymers.

were used to couple the amines of both CHT and PEI. Thefeed molar ratio of PEI to amine of chitosan was 3 : 1 (basedon our preliminary study). Similarly, amidation reactionbetween the carboxyl group on the bifunctional PEG andthe free amine on the grafted CHT-g-PEI through NHS/EDCcross linkers was employed to synthesize the amino terminalCHT-PEI-PEG conjugate for endostatin encapsulation. Thestructural and functional modification of both CHT-g-PEIand CHT-g-PEI-PEG-NH2 conjugates were confirmed byboth 1H NMR and FT-IR spectra as presented in Figures 2and 3, respectively [19–24].

Figures 2(a) and 2(b) show the 1H NMR spectra of thegrafted copolymers and PEG functionalized grafted polymerconjugate, respectively. FromFigure 2(a), characteristic peaksat 𝛿 = 2.9–3.7 ppm were assigned to D-glucosamine unit(H3, H4, H5, and H6) of CHT, the peak at 𝛿 = 3.1 ppm isresponsible for H2, and the peak at 𝛿 = 1.9 ppm is indicativeof the methyl protons of N-acetyl group. The presenceof peaks at 𝛿 = 2.5–3.2 ppm was assigned to methyleneprotons of PEI (–NHCH2CH2–) and confirmed that PEI wassuccessfully grafted onto the CHT chain. Similar results werereported by Sarkar and coworkers [19] for the preparationof fluorescent chitosan-graft-polyethyleneimine and Lui andcoworkers [20] for chitosan-Graft-polyethylenimine/DNAnanoparticles. After the reaction between CHT-g-PEI andCOOH-PEG-NH2, there were new peaks in the 1H spectrumat 𝛿 = 3.6 ppm which belonged to the methylene protons

of PEG (–OCH2CH2–) (relative to Figure 2(a)). There wasa slight downfield shift of the signals of the characteristicprotons of CHT and PEI (Figure 2(b)). This is possibly dueto the deshielding of protons (CHT and PEI) as a result ofaddition of -OCH2 of PEGmoieties. Zhou et al. [24] reportedsimilar downfield shifts for the methylene proton signalsof PEI after conjugation to PEG monomethyl ether with acarboxyl end group (mPEG–COOH).

By comparing the integration of proton signals at 𝛿 =2.5–3.2 ppm and 𝛿 = 2.9–3.7 ppm, the degree of substitutionof PEI per CHTD-glucosamine unit was calculated (see (4)).The integration value of H-2 was defined as 1 during theprocess of calculation.

Degree of substitution (DS) =[𝐼PEI,CH

2CH

2NH] /4

[𝐼CHT,H2,H3−H6] /6

= 3.56/45.51/6= 0.9

(4)

Thereafter degree of grafted (DG) PEI onto CS was calculatedvia

DG = 𝑊2𝑊1× 100%, (5)


O

OH

HONH

4

3

5 2

1

6O

On

HNPEI

-H2, CHT

-H3-H6, CHT

3.68

3.23

3.16 3.00

2.91

2.71

2.66

2.51

4.51 1.003.56

8.5 8.0 7.5 7.0 6.5 6.0 5.5 5.0 4.5 4.0 3.5 3.0 2.5 2.0 1.5 1.0 0.5 0.0(ppm)

–NHCH2CH2 (PEI)

(a)

HONH

OH

O O

O

1

25

3

64

n

HNPEI

-H2, CHT

(PEG)

-H3-H6, CHT

2.29

8.5 8.0 7.5 7.0 6.5 6.0 5.5 5.0 4.5 4.0 3.5 3.0 2.5 2.0 1.5 1.0 0.5 0.0(ppm)

1.000.471.64

–NHCH2CH2 (PEI)

–OCH2CH2 (PEG)

–CH2–NH22.

88

3.29

3.43

3.52

3.61

PEG-NH2

(b)

Figure 2: 1H NMR spectra showing the grafting of polymer conjugates in D2O :CD3COOD (5 : 1). (a) CHT-g-PEI and and (b) CHT-g-PEI-PEG-NH2.

where𝑊1 is mass of pure graft copolymer and𝑊2 is mass ofgraft chains

DG = 24.25214.3

kDa × 100% = 11.3%

𝑊1 = MWCHT + (MWPEI × DS)

= 190 kDa + (25 kDa × 0.97) = 214.3 kDa

𝑊2 = 𝑛 (MWPEI × DS) = (25 kDa × 0.97)

= 24.25 kDa.

(6)

FT-IR spectra of the native and grafted polymers of CHT-g-PEI-PEG-NH2 grafted copolymer are presented in Figure 3.The absorption of ](O-H) and ](N-H) at 3450–3200 cm−1,](C-H) at ∼2930 and 2886 cm−1, ](C = O NH2) at 1656 and1597 cm−1, and ](C-O-C) at 1150 cm−1 in the FT-IR spectrawas assigned to CS D-glucosamine units thereby indicatingthe presence of chitosan as the core polymer framework in thesynthesized CHT-g-PEI-PEG-NH2 conjugate (Figure 3(e)).A new peak appeared at ∼1590 cm−1, which is indicative

of the carboxyl in urea group (Figure 3(c)). Further, theappearance of peaks at 1450.6 and 822.1 cm−1 was attributedto the absorption of -NHCH2CH2- moieties which is fromthe branched PEI, indicating that PEIwas successfully graftedonto CHT [22].

The formation of an amide bond between the carboxylicgroup of PEG and the amine group of CHT-g-PEI copolymerwas noted in the CHT-g-PEI-PEG-NH2 conjugate. As repre-sented in the FT-IR spectrum in Figure 3(e), the characteris-tics peak at 1696.5 cm−1 was assigned to the absorption fromthe C=O stretch in amide bond I, 1463.7 and 1496.4 cm−1were peaks attributed to the absorptions from both the N-Hbend and C-N stretch in amine bond II while the tripartiteabsorption peaks at 1310.0, 1279.6, and 1251.9 cm−1 wereattributed to the N-H bend in plane and C-N stretch in amidebond III. Meanwhile, an absorption peak at 3389.4 cm−1 wasattributed to the N-H stretch from the free primary aminegroup attached to the PEG in the conjugate (CHT-g-PEI-PEG-NH2). Similarly, the characteristic peak of PEI wasrepresented as an absorption peak of 822.1 cm−1 while thepeak at 871.2 cm−1 was assigned to CHT [23]. Notably, the

Journal of Nanomaterials 7%

tran

smitt

ance

Wavenumber (cm−1)1000150020002500300035004000

CHTgPEIPEGNH2ENTCHTgPEIPEGNH2

PEG

CHTPEIPEICHT

(a)

(b)

(c)

(d)

(e)

(f)

Figure 3: FT-IR spectra for native polymers, grafted polymerconjugates, and the endostatin-loaded nanoparticle system.

characteristic peak of PEG at 842 cm−1 contributing to theabsorption of CH2 from –[CH2CH2O]𝑛– was also found inthe spectrum of the CHT-g-PEI-PEG-NH2 polymer conju-gate, which confirms its presence in the polymer construct[24].

The degree of grafting (DG) of PEG onto CS-g-PEI wascalculated based on the 1H NMR using the integration ofproton signals:

DSPEG =[𝐼PEG,

2CH

2NH

2

] /2

[𝐼CHT-g-PEI-PEG,CH2CH

2NH

2

] /4= 0.47/22.29/4

= 0.41

(7)

DG = 0.861215.2

kDa × 100% = 0.40% (8)

𝑊1 = MWCHT-g-PEI + (MWPEG × 0.41)

= 214.3 kDa + (2.1 kDa × 0.41) = 215.29 kDa

𝑊2 = 𝑛 (MWPEG × DS) = (2.1 × 0.41) = 0.86 kDa.

(9)

3.2. Confirmation of the Physicochemical Properties andMorphology of Endostatin-Loaded Nanoparticles. Physico-chemical properties, such as average particle size, sur-face zeta potential, conductivity, and polydispersity index(PDI), of BSA/endostatin-loaded nanoparticles were mea-sured using Dynamic Light Scattering (Zetasizer NanoZS,Melvern Instrument, UK). Table 1 shows the average particlesize for the 15 formulations as influenced by the formula-tion variables. The average size of the nanoparticles rangedbetween 56.28 nm (F11) and 363.10 nm for F14. Although for-mulations with similar concentration of the grafted polymerconjugates (CHT-g-PEI-PEG) seemed to have closely related

average particle size (F2 = 84.22 nm and F4 = 97.38 nm), thesurfactant concentration seemed to be a major determiningfactor for the average particle size distribution among the15 formulations generated using the Box-Behnken design.For instance, the average particle sizes for F5, F6, and F15with a polymer conjugate concentration (1.5mg/mL) were300.10, 295.10, and 310.20 nm, respectively, while F1 and F4with the same polymer conjugate concentration (1.5mg/mL)had average particle sizes of 56.28 and 97.38 nm, respectively,at a ratio of approximately 1 : 3, in comparison to F5, F6,and F15. Both F1 and F4 had a surfactant concentrationof 1% PVA solution while F5, F6, and F15 had surfactantconcentration of 0.55% PVA. Meanwhile, the concentrationof TPP, as a polyionic agent, for the nanoparticle synthesisalso influenced the average particle size of the nanoparticlesamong the 15 formulations. Formulationswith the lowest TPPconcentrations (F1 and F2) (0.1% TPP) had the lowest averageparticle size (F1 = 65.07 nm, F2 = 84.22 nm) compared to F14= 363.10 nm, 0.3% TPP as shown in Table 1. This is becausecrosslinking agents, such as TPP, harden the matrix ofchitosan nanoparticles leading to decreased water absorption[21]. However, higher PVA (1%, 0.55%, 0.55%) and TPPconcentrations (0.3%), as seen in F4, F8, and F11, seemed tobe favorable for the synthesis of moderately sized endostatin-loaded nanoparticles ranging between 56.28 nm for F11 and97.38 nm for F4. In general, increased concentrations ofthe polymer conjugate (2.5mg/mL) and TPP (0.3%) withreduced PVA concentration (0.55%) lead to the synthesis ofa reduced average particle size (F11 = 56.28 nm). This couldbe as a result of strong electrostatic interactions betweenthe BSA/endostatin molecules and the increased positivelycharged amino groups of CHT at higher concentration, asthe main component in the grafted polymer conjugate [25].Furthermore, increased electrostatic interaction between thenegative charge of TPP as a polyanionic agent at higher con-centration and the positively charged amino group of CHTcould possibly account for the reduction in nanoparticle size[25]. Thus, the optimized endostatin-loaded nanoparticleswere prepared at CHT concentration of 2.5mg/mL and havean average hydrodynamic size of 100.6 nm (Figure 4). ThePDI value of the nanoparticles from the 15 formulationsranged between 0.259 and 1.00. Interestingly, nanoparticlesfrom most of the formulations had PDI value less than 1.00showing their uniform disparity in solution. More impor-tantly, the optimized nanoparticles have a PDI of 0.274.

The zeta potential values were positive for most of theformulations (Table 1). Overall, formulations with reducedchitosan concentration in the grafted polymer conjugate(0.5mg/mL) tend to have reduced zeta potential. The zetapotential decreased from amore positive charge towards neg-ative charge surfaces as the polymer conjugate concentrationdecreased from 2.5 to 0.5mg/mL (F3, F7, and F11 comparedto F9, F12, and F13). Meanwhile, as the concentration ofTPP, as the polyanionic agent decreased in the formulations,the PDI values increase across the formulations towardsmore positively charged surfaces with increasing chitosanconcentration in the grafted polymer conjugate (F7 = +11.00compared with F10 = +0.76; and F1 = +16.10 and F2 =+19.80, compared with F4 = −1.87 and F14 = −7.63). It is


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proposed that the overall positive charge on the surface ofthe nanoparticles originates from the protonation of the freeamines on the chitosan component of the grafted polymerconjugate (CHT-g-PEI-PEG-NH2) as chitosan is the mainpolymer in the conjugate. Interactions between the negativelycharged TPP solution and the positively charged chitosan-based polymer conjugate determined the overall surfacecharge of the synthesized nanoparticles. Thus, formulationswith lower TPP concentration and higher concentration ofthe grafted polymer conjugate have more positively chargedsurfaces. Reports have shown that the surface of tumor cellsis negatively charged [26–28]. Interestingly, the optimizedendostatin-loaded nanoparticle has an average positive zetapotential value of 7.95mV (Figure 4) which predisposed itfor direct targeting in squamous cell carcinoma, as laterpresented in this study [29].

Both TEM and SEMmicrographs (Figures 5 and 6, resp.)confirmed the morphology of endostatin-loaded nanoparti-cles to be spherical with transparent surfaces. Nanoparticleshape and size have been reported to influence internal cellu-lar binding and uptake of drug-loaded nanoparticles [30–32].Importantly, spherical shape nanoparticles possess enhancedcellular binding and internalization compared to nanoparti-cles with other shape configurations [33]. The TEM imagesfurther confirmed that endostatin-loaded nanoparticles withan average size less than 100 nmwere successfully synthesizedin agreement with the DLS average size measurement. TheSEM micrograph also suggested that some of the smallerparticles possibly agglomerate to form larger nanoparticles[34] ranging between 200 and 400 nm. However, it should benoted that SEM images of the nanoparticles are acquired intheir dried solid state as opposed to TEM and DLS averagesize measurements where the nanoparticles were dispersedin liquid. Interestingly, the average size of the optimized

Figure 5: TEM images of optimized endostatin-loaded nanoparti-cles.

Table 2: Drug-loading capacity and entrapment efficiency.

Formulations Entrapment efficiency Loading capacity (%)F1 0.943 31.43F2 0.953 31.77F3 0.975 32.50F4 0.928 30.93F5 0.984 32.80F6 0.980 32.67F7 0.947 31.57F8 0.973 32.43F9 0.959 31.97F10 0.966 32.20F11 0.963 32.10F12 0.969 32.30F13 0.981 32.70F14 0.971 32.37F15 0.964 32.13

endostatin-loaded nanoparticles are within the range forsubcutaneous delivery intended for this study [35, 36].

3.3. Confirmation of BSA/Endostatin Loading Capacity andEntrapment Efficiency of Nanoparticles. As presented inTable 2, the drug-loading capacity and entrapment efficiencyof the 15 formulations showed that BSA, as a model drug, wassuccessfully loaded and entrapped within the grafted CHT-based polymeric network and ranged between 92.8% (F4)and 98.4% (F5). This could be attributed to the influenceof PEG coating on the surface of the nanoparticle. PEGcoating has been found to increase encapsulation efficiencyas free dissolved drugs are captured within the matrix ofnanoparticles during PEGylation [21]. In general, formula-tions with lower CHT concentrations (0.5mg/mL) showedhigher loading and entrapment capacity (F8, F12, and F13compared to F7, F10, and F11 with CHT concentration of2.5mg/mL, Table 1). This could possibly be due to the lowerviscosity that increases encapsulation of BSA or endostatinand possible promotion of gelation between CHT and TPP


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Figure 6: Gold-platinum sputtered SEM micrographs of the optimized ENT-loaded nanoparticles. (a) Magnification = 8.36Kx. (b)Magnification = 50.00Kx.

as the cross linking agent [21].This corroborates the report ofVandenberg et al. [37] that the viscous nature of the gelationmedium affects the drug encapsulation during the synthesisof chitosan microspheres. More so, effective confinementof BSA/endostatin molecules in the nanoparticles could beattributed to the electrostatic interaction between the pro-teins and the long chain of chitosan. Formulations F8, F11, andF14 with higher TPP concentrations (0.3%) showed slightlyhigher encapsulation. This could be a result of increasedcrosslink density which possibly prevents the leaching awayof entrapped BSA/endostatin molecules during formation ofnanoparticles [21]. Further, high TPP concentration with afixed chitosan concentration may increase the solution pHwhich consequently influences the overall negative chargeat the surface of the BSA molecules. This enhances theelectrostatic interactions between chitosan chains and BSAmolecules and thus results in higher BSA loading andencapsulation efficiency.

3.4. Nanoparticle Swelling Behaviour and In Vitro DrugRelease. As presented in Figure 7, the swelling capacity ofBSA-loaded nanoparticles increased with time at the varyingCHT concentrations. The degree of swelling was greater atpH 6.8 than at pH 7.4, which would ultimately contributeto an enhanced release of the bioactivity at tumoral pH.The effect of PEI and the surface shielding effect of PEGcontributed to the swelling behaviour and BSA/endostatinrelease from the nanoparticles [21]. Overall, nanoparticleswelling reached equilibrium after 2 hours in the buffersystem. RP-HPLC, as presented in Figure 8, was employedto quantitate the in vitro BSA/endostatin release from theBSA/endostatin-loaded nanoparticles as shown in Figure 11.BSA elute at 2.216 minutes over a run time of 6 minutes andresult corroborates to previous report of Mukhopadhyay [38]and Mukhopadhyay et al. [39].

3.5. Effects of Chitosan and TPP Concentrations on Swellingand Release Kinetics. Chitosan concentration has beenreported to play a major role in determining the swellingcapacity and drug release kinetics of CHT-based nanopar-ticles [12]. Formulations with higher CHT concentrations

showed a higher swelling ratio with increased BSA releasefrom the nanoparticles. Furthermore, the concentration ofthe cross linker (TPP) has also been reported to influence theswelling behaviour of CHT-based nanoparticles [21]. LowerTPP concentrations allow for weak electrostatic interactionwith the protonated amines of CHT and thus facilitate massmovement of the buffer solution into the matrix of thenanoparticles. At acidic pH (6.8 as presented in this study),the mass ratio of available protonated amino groups of CHTwhich can interact with TPP anions increased at a constantor lower TPP concentration (Figure 7(a)).This in turn resultsin a more loosely packed nanoparticle matrix which canbe easily hydrated by the buffer solution and facilitate therelease of the entrapped BSA/endostatin into the system. Forinstance, F10 has the highest cumulative BSA release at bothpH 6.8 (66%) and pH 7.4 (58%) at CHT concentration of2.5mg/mL and TPP concentration of 0.2%, compared to F8having cumulative BSA release of 48% at pH6.8 and 40%BSArelease at pH 7.4 at CHT/TPP concentrations of 0.5mg/mLand 0.3%, respectively (Figure 9).

3.6. Effects of pHonNanoparticle Swelling andBSA/EndostatinRelease. The tumor microenvironment, including squamouscell carcinoma, and normal cellular physiological conditionwere both mimicked at pH 6.8 and 7.4, respectively [40],to investigate the swelling behaviour and in vitro releasekinetics of BSA-loaded nanoparticles. Overall, BSA releasefrom within the nanoparticle matrix was increased at acidicpH 6.8 compared to physiological pH of 7.4 across allformulations over a period of 16 hours (Figure 9(b)). SinceCHT is the main component of the grafted polymer networkforming the nanoparticles, the pH-responsive characteristicsof the BSA-loaded nanoparticles should be linked to thehydrolysis of CHT at changing pH conditions.The free aminegroups in the CHT moiety within the polymer conjugatebecome protonated at acidic pH, forming positively chargedNH3+ groups which are hydrophilic. This further results

in an electrostatic repulsion among the protonated aminogroups which weaken both the intra- and intermolecularhydrogen bonding interaction of CHT molecules. This inturn allows the enhanced permeation of the buffer solution


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Figure 8: Reverse Phase HPLC chromatogram of BSA. Retention time of BSA as a model for endostatin is 2.216 minutes and run time is6 minutes. No interference was observed for RP-HPLC for blank sample. Elusion was monitored at wavelength of 280 nm to detect BSA aspreviously reported [38].


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Figure 9: Cumulative drug release of the 15 experimental design formulations at pH 6.8 (a) and 7.4 (b). BSA release is higher at acidic pH of6.8 than at physiological pH of 7.4 for each corresponding formulation.


into the network therefore causing the equilibrium swellingratio to increase as opposed to what is obtained at bothneutral and basic pH conditions. In this study, swelling andshrinking mechanism with regard to swinging pH values hasbeen investigated in terms of CHT-based smart responsivenanoparticulate systems for localized endostatin delivery.

3.7. Effects of PEI and PEGylation on Swelling and DrugRelease. The protonating effects of the amines in PEI atlow pH value contributed to the increased swelling of thenanoparticles as presented in Figure 7(c). At low pH, theamino groups on both CHT and PEI become protonatedand thereby increase the overall NH3

+ groups, which arehydrophilic. Thus, the buffering system could easily gainaccess into the nanoparticle network which facilitates anincrease in the equilibrium swelling ratio. Meanwhile, theamphiphilic nature of PEG in the grafted polymer networkpossibly influenced the swelling property as well as the releasepattern from the nanoparticles. PEGylation has increasedthe circulation time of PEG-modified nanosystems as wellas prolonging the release of encapsulated drug in nanopar-ticulate systems [41–43]. In our system, the carboxyl groupin the bifunctional PEG was used in grafting the aminegroup of PEI moiety in the CHT-based polymer conjugate.PEG is amphiphilic; its hydrophilic property depends onwhich part of the polymer is used during covalent reaction.The sheathing effect of PEG on the nanoparticle surfacetends to increase the compact nature of the nanoparticulatesystem and, thus, reduce the swelling ratio of the drug-loaded nanoparticle which in turn prolongs the release ofthe encapsulated drugs within the nanoparticle matrix. Asevidenced in Figure 7(c), CHT-g-PEI-PEG nanoparticleshave a reduced hydrated size as compared to both CHT-PEI and CHT nanoparticle. At equilibrium swelling pointafter 2 hours, CHT-g-PEI-PEG-NH2 nanoparticles had anaverage size of 458.9 nm, while both CHT-g-PEI and CHTnanoparticles had sizes of 794.2 and 534.2 nm, respectively.

3.8. Optimization of the Experimentally Derived Box-BehnkenFormulations. The results of the response optimization pro-cedure (MINITAB, V15, Minitab, USA) showed a direct rela-tionship between the fitted data and the observed response.The surface plot results (Figure 10) showed the effects of thepolymer complex concentration, the surfactant concentra-tion, and TPP concentration on the overall average size ofnanoparticles and their surface charge as well as ENT releasefrom the nanoparticle matrix at changing pH conditions of6.8 and 7.4. Nanoparticle size and ENT release at the loweracidic pH of the tumor microenvironment were maximizedwhile zeta surface potential and ENT release at physiologicalpH were minimized.These responses were chosen to achievean optimum desirability in terms of the nanoparticulatesystem performance in cancer nanomedicines. The opti-mized levels of the independent variables, the correspondingpredicted response, 𝑦, and both the individual and com-posite desirability values are presented in Figure 11. Theoptimized formulation variables were 1.22mg/mL for thegrafted polymer conjugate (1.22mg/mL CHT, 0.25% PEI,and 1mg/mL functionalized PEG), 0.10% PVA, and 0.12%

TPP for the synthesis of endostatin-loaded nanoparticles.Based on the statistical desirability function, it was foundthat the composite desirability of the formulation was 0.86.The experimentally derived values for the nanoparticle size,surface zeta potential, and ENT in vitro release of theoptimized formulation was in close agreement with thepredicted values (size = 100.6 nm, PDI = 0.274, surface zetapotential = 7.95mV, and 𝑡8 h = 0.39), demonstrating thereliability of the optimization procedure in predicting theunique behaviour of the desired nanoformulation having anaverage nanoparticulate size and surface charge for efficientcellular binding and uptake and internalization as well ashaving an improved, sustained, and efficient release patternof ENT at tumor acidic microenvironment, as demonstratedin Figure 12.

3.9. Confirmation of Cytocompatibility and AntiangiogenicEffect of Endostatin-Loaded Nanoparticles. Cytotoxicity ofendostatin-loaded and unloaded nanoparticles and theirantiangiogenic effects on oesophageal squamous cell car-cinoma (OSCC) were assessed by treating KYSE-30 cellline with varying concentrations of these nanoformulations.There is no significant difference in cell viability of the nativepolymer conjugate when compared to the control experiment(untreated cells) at polymer concentrations ranging between125 and 500𝜇g/mL (Figure 13). Percentage cell viability wasabove 80%. A major significant difference only occurredat a concentration of 1000 𝜇g/mL of the polymer conjugate(63.85% of native polymers conjugate treatment when com-pared to the control). Thus, this formulation has potentialfor use in the delivery of cancer therapeutics without toxicityconcern since percentage cell viability of 80% or above showsgood cytocompatibility of tested formulations [44, 45].

Although the polymers employed for nanoparticle syn-thesis were shown to have good cytocompatibility and tobe not toxic, it is possible for nanoparticles to displayspecific toxicity because of their size and their accumulationand degradation in cells after internalization or due to theaddition of other excipients during their preparation [46].Our findings indicate that blank nanoparticles from thepolymer will not affect the results from the proliferationassay performed with endostatin-loaded nanoparticles forthe same treatment time. As such, oesophageal squamous cellcarcinoma (KYSE-30) cells’ proliferation was subsequentlyexamined in response to endostatin-loaded nanoparticles.

Reports have shown that endostatin blocks endothelialcycle progression and reduces the expression of proliferationgenes [47]. It induces endothelial cell proliferation [48] andapoptosis pathways in the 𝜇g/mL range [49]. The antiangio-genic efficacy of endostatin-loaded nanoparticles has beenwell reported [10, 50, 51]. As presented in our study (Fig-ure 13), endostatin-loaded nanoparticles showed a significantreduction in cell viability (38.32%) over a concentration of1000 𝜇g/mL. We observed 5.32% proliferation inhibition inresponse to nanoparticle concentration of 125 𝜇g/mL over24 hours’ incubation period. Meanwhile, the proliferationinhibitory effect doubled to 13.36% at 250 𝜇g/mL endostatin-loaded nanoparticle concentration over the same treatmentperiod. Interestingly, a dramatic increment in proliferation


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inhibition, more than fourfold (61.68%), was observed atnanoparticle concentration of 1000𝜇g/mL over the sameperiod of incubation. An unpaired t-test and two-tailed 𝑃values for native and endostatin-loaded nanoparticles whencompared to the control untreated cells showed statisticallysignificant differences at 𝑃 = 0.0388 and 𝑃 = 0.0003, respec-tively. Meanwhile, there is no significant difference betweennative and endostatin-loaded nanoparticles for 125, 250, and500 𝜇g/mL after 24 hours’ treatments when compared to thecontrol. However, longer exposure of KYSE-30 cells to longerincubation time may facilitate the inhibition potential of thedrug-loaded nanoparticles relative to the native particles whencompared to the control.

On a general note, encapsulation with nanocarriersreduces the toxicity of drug as they offer protection againstdegradation by RES [52, 53]. In addition, encapsulation couldalso enable nanoparticles to target cell membranes by ensur-ing the release of drug close to the cell surface and increase thecellular uptake of nanoparticles in specific targeting studieswhich could enhance site-specific delivery of the nanocarriersin targeted cells thereby increasing their efficacy [54]. Itis interesting to note that the doses of endostatin-loadednanoparticles presented in this study, administered to theKYSE-30 cells to achieve the antiangiogenic drug levels pre-sented in Figure 12, were in the range of 125–1000 𝜇g/mL andwere not expected to affect cell proliferation. The positively


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charged surface of the designed nanocargo, as presentedin Figure 4(b), could have facilitated direct binding to thenegatively charged surface of KYSE-30 cells having a typicaltumor cell surface [26]. Similarly, the shielding effect of thenanoparticles’ surface by PEG could have modulated therelease of endostatin within the nanoparticle matrix therebyincreasing its antiangiogenic efficacy on the treated cells.

Formation of new blood vessels and cell proliferation arepotent indicators of angiogenesis [55, 56] which are key fortumor cell survival. Therefore, our novel nanosystem couldserve as a potential nanocargo for the delivery of endostatinand other related antiangiogenic cancer therapeutics foreffective management of SCC.

4. Conclusions

The study confirmed successful conjugation, formulation,and optimization of a PEGylated CHT-g-PEI-based nanosys-tem loaded with endostatin as a potential delivery vehiclefor endostatin. The concentration of CHT as the mainpolymer as well as the cross linker concentration wasa fundamental factor that modulates the BSA/endostatinencapsulation, release, and swelling behaviour. More so,addition of PEI onto the CHT moiety increased the overallamine protonation of the grafted polymer network at lowerpH value and thus impacted positively on nanoparticleswelling and in vitro BSA/endostatin release. Positivelycharged spherical nanoparticles were synthesized with avarying BSA/endostatin release profile at tumor pH andnormal physiological pH conditions were observed, dueto instantaneous swelling and shrinking behaviour of thenanosystem, respectively. The positively charged surface ofthese nanoparticles could be harnessed for adhesion to thenegatively charged surface of many tumor cells includingOSCC. The polymer conjugate was validated to be nontoxicto established squamous cell oesophageal carcinoma cellline (KYSE-30). Furthermore, encapsulation of endostatin

Untreated cellsNative nanoparticlesEndostatin-loaded nanoparticles

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Figure 13: Cell viability/proliferation of treated and untreatedoesophageal squamous cell carcinoma (KYSE-30) after 24-hourexposure to the nanoformulations at varying polymer concen-trations. An unpaired 𝑡-test, two-tailed 𝑃 values for native andendostatin-loaded nanoparticles when compared to the controluntreated cells showed statistically significant differences at 𝑃 =0.0388 and 𝑃 = 0.0003, respectively.Meanwhile, there is no significantdifference between native and endostatin-loaded nanoparticles for 125,250, and 500 𝜇g/mL after 24-hour treatment.

in the nanoconstruct showed efficient cytotoxicity againstcell proliferation and angiogenesis in a tumor model invitro. This study validates the potential use of endostatin-loaded CHT-g-PEI-PEG-NH2 nanoparticles in OSCC invitro; further in vivo experiments are necessary to confirmthe preclinical potential of this system. Further, function-alizing this nanoconstruct with OSCC specific biomarkerscould enhance direct targeting of endostatin as a potentendogenous antiangiogenic inhibitor in OSCC managementboth in vitro and in vivo.

Conflicts of Interest

The authors confirm that there are no conflicts of interest.

Acknowledgments

This work was funded by the National Research Foundation(NRF) of South Africa and the Cancer Research Associationof South Africa (CANSA).

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[4] V. P. Torchilin, “Targeted pharmaceutical nanocarriers forcancer therapy and imaging,” The AAPS Journal, vol. 9, no. 2,pp. E128–E147, 2007.

[5] J. Folkman, “Tumor angiogenesis: therapeutic implications,”The New England Journal of Medicine, vol. 285, no. 21, pp. 1182–1186, 1971.

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