AlginateâC18 Conjugate Nanoparticles Loaded in Tripolyphosphate

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Alginate-C18 conjugate nanoparticles loaded in tripolyphosphate-crosslinked chitosan-oleic acid conjugatecoated calcium alginate beads as oral insulin carrier Mulham Alfatama, Lee Yong Lim, and Tin Wui Wong Mol. Pharmaceutics, Just Accepted Manuscript • DOI: 10.1021/acs.molpharmaceut.8b00391 • Publication Date (Web): 11 Jul 2018 Downloaded from http://pubs.acs.org on July 13, 2018

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Molecular Pharmaceutics

1

Alginate-C18 conjugate nanoparticles loaded in tripolyphosphate-crosslinked

2

chitosan-oleic acid conjugate-coated calcium alginate beads as oral insulin

3

carrier

4

Mulham Alfatama,a,b Lee Yong Lim,c Tin Wui Wonga,b

5

a

6

b

7

Universiti Teknologi MARA Selangor, Puncak Alam, 42300, Selangor, Malaysia.

8

c

9

University of Western Australia, 35 Stirling Highway, Crawley WA 6009, Australia.

Non-Destructive Biomedical and Pharmaceutical Research Centre, iPROMISE Particle Design Research Group, Faculty of Pharmacy

Pharmacy, Centre for Optimisation of Medicines, School of Allied Health, The

10 11

Running title: Alginate nanoparticles-in-beads as oral insulin carrier

12 13

*Corresponding

author.

Tel.:

+60

3

32584691.

E-mail

14

[email protected], [email protected] (T.W. Wong).

addresses:

15 16

Abstract. Simple alginate, alginate-stearic acid and alginate-C18 conjugate

17

nanoparticles, and tripolyphosphate-crosslinked chitosan-oleic acid conjugate-coated

18

calcium alginate beads as the vehicle of nanoparticles were designed. Their size, zeta

19

potential, morphology, drug load, drug release, matrix molecular characteristics,

20

mucus penetration, HT-29 cell line cytotoxicity and intracellular trafficking, in vivo

21

blood glucose lowering and insulin delivery profiles were characterized. Alginate-C18

22

conjugate nanoparticles were non-toxic. Among all nanoparticles variants, they had

23

reduced size and zeta potential thus enhancing particulate mucus penetration and

24

intracellular trafficking. Their insulin reabsorption tendency was minimized as

25

alginate active COOH/COO- sites were preoccupied with C18. Their loading into

26

coated beads was translated to reduced drug release in simulated gastric phase with

27

nanoparticles being released in the intestinal phase. The combination dosage form

28

increased the blood glucose lowering extent of insulin, and blood insulin level instead

29

of nanoparticles or beads alone. Nanoparticles in beads represented a viable approach

30

in oral insulin delivery.

31 32

Keywords: Alginate; Beads; Chitosan; Insulin; Nanoparticles; Oral 1 ACS Paragon Plus Environment

Molecular Pharmaceutics 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

1 2

Introduction

3 4

Diabetes mellitus is an endocrine disorder in which the process of

5

carbohydrate metabolism is disturbed due to deficiency in insulin secretion, insulin

6

resistance or both.1-2 Often, the diabetic patients require insulin injections to maintain

7

normal blood glucose levels and minimize long-term medical complications, such as

8

organ degeneration.3 Injections are invasive, and the deterioration in quality of life

9

caused by the dermal trauma and pain of daily insulin injections frequently leads to

10

poor patient compliance.4-5 For this reason, there has been intense research over the

11

past few decades in the development of non-injectable insulin therapy.6-7 The peroral

12

route has emerged as an attractive non-invasive route for insulin administration as it

13

mimicks the physiological fate of endogenous insulin, which may provide for a better

14

glucose homeostasis.8-10

15

Delivery of insulin by the oral route is, however, associated with significant

16

challenges, namely insulin degradation by the gastrointestinal (GI) enzymes, and the

17

presence of insulin absorption barrier ascribed to epithelial mucosa, mucus lining and

18

tight junction.11-12 To overcome the absorption barrier, nanocarriers have been

19

advocated to promote mucoadhesion and prolongation of GI residence time to allow

20

for the transmucosal penetration of carrier and/or insulin.6,13-17 Nonetheless, the use of

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nanocarriers as oral insulin delivery systems has translated to low insulin

22

bioavailability due to rapid premature release and enzymatic digestion of the insulin

23

load.18-19

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Alginate is a polysaccharide derived from the brown seaweeds, such as

25

Laminaria spp., Ascophyllum spp., and Sargassum spp..20 It is an unbranched

26

copolymer consisting of alternating blocks of 1-4 linked α-L-guluronic acid (G-block)

27

and β-D-mannuronic acid (M-block) residues. As a pH-responsive polymer, alginate

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dissolves at near neutral and alkaline pH but exists as water-insoluble alginic acid at

29

low pH. Dissolved alginate is capable of forming a gel through Na+/Ca2+ exchange

30

and the formation of ionic crosslinks with the Ca2+ via its carboxylate moiety. Calcium

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alginate-insulin nanoparticles have been developed,15,21 and the administration of

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these nanoparticles by gastric gavage is found to successfully reduce the blood

33

glucose level of diabetic rats. Besides providing insulin systemically, the nanoparticles

2 ACS Paragon Plus Environment

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may also reduce intestinal glucose absorption via alginate-glucose interaction in the

2

intestinal lumen.

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Previous studies have suggested that the alginate gel is characterized by a high

4

level of porosity and rapid drug release profiles.22-23 Thus, the successful application

5

of alginate-insulin nanoparticles via gastric gavage may not necessarily translate to an

6

equivalent therapeutic outcome had the nanoparticles been administered perorally. On

7

this note, the present investigation aimed to microencapsulate the alginate-insulin

8

nanoparticles into alginate beads coated with an oleic acid-conjugated chitosan layer.

9

The purpose of introducing the coated bead as a vehicle for the alginate-insulin

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nanoparticles is multiple-fold. Firstly, the beads may serve to better protect the insulin

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from premature release into the gastric medium. Secondly, the alginate upon

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ionization and dissolution in the intestinal medium is mucoadhesive and may further

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prolong the intestinal residence time of the nanoparticles. Lastly, the beads can

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provide additional alginate material for the binding of intestinal glucose thus further

15

reducing glucose absorption into the blood stream. The summative effect was deemed

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to lead to a more definitive reduced hyperglycemia. The chitosan-oleic acid conjugate

17

was used as the coating material instead of chitosan. The chitosan is soluble in acidic

18

medium and is prone to dissolve in the gastric milieu. Its coacervation with the

19

alginate matrix of beads may be lost upon oral administration. Grafting of

20

hydrophobic oleic acid onto chitosan is envisaged to reduce the dissolution propensity

21

of chitosan, and promote its interaction with alginate as a water-insoluble coacervate.

22

Alginate is known to interact with insulin to form insoluble complexes15 that

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then lowers insulin bioavailability. In utilising alginate not only in the nanocarrier but

24

also in the supportive bead structure, there was a risk of excessive alginate

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complexation that would negate the anti-diabetic action of the insulin load. To manage

26

the complexation reaction, this study designed three variants of alginate

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nanoparticulate carriers for the insulin load: simple alginate nanoparticles, alginate-

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stearic acid nanoparticles and alginate-C18 nanoparticles. Stearic acid is a saturated

29

fatty acid with a C18 chain length, and it was physically embedded within the alginate

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matrix to prepare the alginate-stearic acid nanoparticles. Alginate-C18 nanoparticles,

31

on the other hand, were prepared using alginate covalently grafted with C18 aliphatic

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chains. In both types of nanoparticles, the hydrophobization of alginate was

33

hypothesised to enable a polymer-drug interaction that conferred delayed insulin

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release characteristics without curtailing insulin bioavailability. The present study 3 ACS Paragon Plus Environment


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aims to investigate and identify the potential nanoparticles in beads as oral insulin

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carrier.

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2. Experimental section

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Materials. Bovine insulin (Sigma Aldrich, USA) was employed as the drug of

7

interest. Guluronic acid-rich sodium alginate (molecular weight = 4.8 × 106 ± 1.8 ×105

8

Da, nominal viscosity = 300 mPa.s, mannuronate/guluronate ratio = 0.59; Manugel®

9

DMB, ISP, USA), chitosan (molecular weight = 1.5 × 106 Da, degree of deacetylation

10

= 86 %; Zhejiang Aoxing Co. Ltd., China), stearic acid (Hesego Industry Sdn Bhd,

11

Malaysia), 1-bromooctadecane (Sigma Aldrich, China) and oleic acid (Merck,

12

Germany) were used as received. Acetonitrile (Merck, Germany), trifluoroacetic acid

13

(BDH, UK), and ultrapure water processed at 18 MΩ were utilized to prepare the

14

mobile phase for high performance liquid chromatographic (HPLC) analysis.

15

Streptozotocin (Sigma Aldrich, USA) was employed for induction of diabetes in rats,

16

while ketamine hydrochloride and xylazine hydrochloride (Troy Laboratories,

17

Australia) were used as anaesthetic agents in the animal study. All other materials

18

were of analytical grade. Distilled water (ELGA, UK) was used throughout.

19 20

Synthesis

of

alginate-C18

conjugate.

Alginate-C18

conjugate

was

21

synthesised using the protocol described by Valle, & Romeo (1990).24 Briefly, sodium

22

alginate (2% w/w in water, 100 ml) was transformed into alginic acid by adjusting the

23

solution pH to 2.5 using 0.3 M hydrochloric acid (HCl). After stirring for 12 h at 250

24

rpm and 25°C, the gel suspension was soaked in absolute ethanol (50 ml), stirred for 5

25

min and filtered. The alginic acid (1 g) was dissolved in 15 ml of tetrabutylammonium

26

hydroxide (TBA; Sigma-Aldrich, Switzerland), and lyophilized (Eyela, Japan). The

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lyophilized TBA-alginate (1 g) was dissolved in 100 ml of dimethylsulfoxide

28

overnight, then reacted for 24 h with 1-bromoctadecane (0.2850 g) under continuous

29

magnetic stirring at 250 rpm and 25ºC. Sodium chloride (96 ml, 2.5 M in water) was

30

added to transform the alginate-C18 into a salt. A strong gel was formed by adding 30

31

ml of 70 % ethanol, and after agitation for 15 min at 250 rpm, the alginate-C18 gel

32

was washed in two separate runs with 250 ml of 70 % ethanol and 50 ml of acetone.

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The gel was dried in a hot air oven at 25°C for 24 h followed by drying under silica

34

gel in a desiccator.25 4 ACS Paragon Plus Environment

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1 2

Synthesis of fluorescein isothiocyanate (FITC)-conjugated alginates.

3

Alginate-based 1,3-diaminopropane conjugates were prepared by activating 100 g of

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1% w/w alginate or alginate-C18 conjugate in the presence of 80 mg of 1-ethyl-3-(3-

5

dimethylaminopropyl) carbodiimide (EDC) and 60 mg of N-hydroxysuccinimide

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(NHS)26 at 250 rpm at 25ºC for 4 h. 1,3-diaminopropane at 1 ml was slowly added and

7

the reaction allowed to proceed for 24 h. The resultant product was isolated by

8

precipitation with 100 ml acetone, washed with 100 ml ethanol, dried in the fumehood

9

over 2 h, and further dried in the oven at 40°C for 24 h.

10

The

alginate-1,3-diaminopropane

or

alginate-C18-1,3-diaminopropane

11

conjugate (500 mg in 100 ml of water) was reacted with FITC (5 mg in 25 ml of

12

methanol) at 1000 rpm over 24 h in the dark at 25ºC. The FITC-conjugated alginates

13

were precipitated by adding 300 ml of acetone, then recovered by centrifugation at

14

3400 rpm for 5 min at 25°C. The products were purified by dialysis (molecular weight

15

cut-off: 1000 Da; 6 Spectra/Por, Spectrum Laboratories, USA) against water for 24 h

16

in the dark, and freeze dried in the dark. Samples were stored under silica gel in a

17

desiccator.

18 19

Synthesis of chitosan-oleic acid conjugate. Oleic acid was covalently

20

conjugated to chitosan via an amide linkage through an EDC-mediated reaction.27

21

Chitosan (1 g), dissolved in 100 ml of 1%v/v acetic acid and further diluted with 100

22

ml of methanol (Merck, Germany), was mixed with oleic acid (0.7 g) followed by

23

dropwise addition of 15 ml of EDC (0.07 g/l in methanol) under continuous magnetic

24

stirring at 250 rpm at 25ºC. After 24 h of reaction, the chitosan-oleic acid product was

25

precipitated by transferring the reaction mixture into 200 ml of 7:3 v/v methanol:25 %

26

aqueous ammonia (Merck, Germany). The chitosan-oleic acid was collected by

27

filtration, washed sequentially with 50 ml each of methanol, diethyl ether (Merck,

28

India) and distilled water, and dried in the oven at 40°C for 48 h.

29 30

Preparation of nanoparticles. Three types of insulin-loaded alginate

31

nanoparticles were prepared and they were abbreviated as simple alginate

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nanoparticles (AN), alginate-stearic acid nanoparticles (ASAN), and alginate-C18

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nanoparticles (AC18N). AN and AC18N were prepared by first dropwise addition of

34

insulin (5 mg dissolved in 40 ml of 0.01 M HCl) to an alginate solution (100 mg of 5 ACS Paragon Plus Environment


1

sodium alginate or sodium alginate-C18 conjugate in 60 ml of 0.01 M NaOH) under

2

continuous magnetic stirring at 1000 rpm at 25°C. The pH of the mixture was 4-4.5.

3

To prepare the ASAN, stearic acid (10 mg in 10 ml ethanol) was first dispersed into

4

the alginate solution prior to the addition of insulin through magnetic stirring at 500

5

rpm. The resultant mixtures, under continuous magnetic stirring, were subsequently

6

spray dried (TwinNanoSpray, UiTM, Malaysia) under the following operating

7

parameters: inlet temperature = 60°C, outlet temperature = 23 ± 2°C, solution feed

8

rate = 4 ml/min, concurrent air flow rate = 2-2.5 m/s, atomizing air pressure = 6 bar.

9

The spray dried powders were retrieved into 10-ml amber diagnostic vials and stored

10

at 25°C under silica gel in a desiccator.

11 12

Preparation of nanoparticles-loaded beads. Insulin-loaded nanoparticles

13

(70.8 mg) were dispersed into a sodium alginate solution (10 g, 2.5 %w/w in water) by

14

magnetic stirring for 10 min at 250 rpm. The dispersion was processed into beads by

15

means of an encapsulator equipped with a vibrating nozzle device (Nisco,

16

Switzerland) using the following processing parameters: dispersion flow rate = 0.5

17

ml/min, nozzle vibrational frequency = 6 KHz, coaxial nozzle with internal diameters

18

of 600 µm and 400 µm respectively, processing temperature = 25ºC. Droplets of

19

dispersion extruded from the encapsulator were crosslinked in 50 ml of 4 %w/w

20

calcium acetate solution under continuous magnetic stirring at 750 rpm. The distance

21

between the vibrating nozzle and the surface of the crosslinking solution was fixed at

22

25 cm. The crosslinking reaction was continued under stirring for 10 min after the

23

addition of the last drop of dispersion. The formed beads were harvested, washed

24

thrice with 15 ml of distilled water and oven-dried (Memmert, Germany) at 40°C over

25

24 h.

26

To coat the beads, a 0.5 %w/w chitosan-oleic acid conjugate in 0.5 %v/v acetic

27

acid was prepared and its pH was adjusted to 5.5 with 2 M HCl or NaOH, to promote

28

the coacervation of alginate (pKa = 3.65) with chitosan (pKa = 6.2-7) at the interface

29

of the alginate beads. Sodium tripolyphosphate, at 1 %w/w expressed with respect to

30

the weight of alginate, was added into the nanoparticulate dispersion. Ten ml of

31

nanoparticulate dispersion and 1.8 ml of chitosan-oleic acid conjugate solution were

32

introduced concurrently at 0.5 ml/min and 0.09 ml/min respectively to the nozzle head

33

of the encapsulator. The nanoparticulate dispersion emerged from the nozzle

34

simultaneously with the chitosan-oleic acid conjugate solution, which enabled the 6 ACS Paragon Plus Environment

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1

coacervation of chitosan-alginate to occur at the bead interface before the beads fell

2

into the calcium acetate solution (50 ml, 4 %w/w). Further consolidation of bead

3

matrix and coating occurred when the tripolyphosphate ions crosslinked with chitosan,

4

and the calcium ions reacted with alginate when the coated beads were stirred in the

5

calcium acetate solution for a further 10 min. The coated beads were processed using

6

the same conditions as the uncoated beads unless otherwise stated.

7

Nuclear magnetic resonance (NMR) spectroscopy. High resolution 1H-

8

NMR spectra of alginate, 1-bromooctadecane and alginate-C18 conjugate, as well as,

9

chitosan, oleic acid and chitosan-oleic acid conjugate were recorded using the

10

UltrashieldPlus 500 nuclear magnetic resonance spectrometer (Bruker, Germany) at

11

25ºC. Prior to 1H-NMR analysis, the alginate and alginate-C18 conjugate were

12

subjected to partial acid hydrolysis in order to obtain well resolved signals. Briefly,

13

0.1 g of sample dissolved in 100 ml of water was adjusted to pH 5.6 by adding 0.1 M

14

HCl, and the solution was heated at 100°C for 1 h, reduced to pH 3.8 and heated at

15

100°C for another 30 min. The solution was cooled to room temperature, neutralized

16

to pH 7-8 with 0.1 M sodium hydroxide (NaOH) and freeze-dried. The dried sample

17

(10 mg) was dissolved in 0.5 ml deuterium oxide (D2O).28 Chitosan and chitosan-oleic

18

acid conjugate were dissolved at 1 %w/w in deuterated water acidified with HCl to pH

19

4.0.29 The solutions were frozen and thawed in three repeated cycles to exchange the

20

labile proton of chitosan and oleic acid with deuterium. 1-bromooctadecane and oleic

21

acid were dissolved in deuterated DMSO (Merck, Germany). Triplicates were

22

conducted.

23 24

Fourier transform infrared (FTIR) spectroscopy. Two mg of sample

25

together with 78 mg of potassium bromide (FTIR grade, Aldrich, Germany) were

26

compressed into a disc for FTIR analysis at a resolution of 4 cm−1 over a wavenumber

27

region of 450-4000 cm−1 (Spectrum RX1 FTIR system, Perkin Elmer, USA). The

28

characteristic peaks of IR transmission spectra were recorded. At least triplicates were

29

carried out for each batch of sample and the results averaged.

30 31

Particle size and zeta potential. The particle size and zeta potential of

32

nanoparticles dispersed in 96 % ethanol through brief sonication (30 s) were measured

33

by means of photon correlation spectroscopy technique (Malvern Zetasizer Nano ZS

34

90, Malvern Instruments Ltd., UK) at 25°C in quartz cell and zeta potential cell 7 ACS Paragon Plus Environment


Page 8 of 43

1

respectively at a detection angle of 90°. Triplicates experiments were conducted and

2

the results averaged.

3 4

Surface morphology. The morphology of nanoparticles was examined using

5

transmission electron microscopy (Tecnai G2 20S TWIN, FEI, The Netherland). A

6

drop of nanoparticles dispersed in acetone (Fisher Scientific, UK) was placed on a

7

carbon-coated copper grid of 200 mesh, and allowed to dry at 25°C prior to

8

microscopic viewing at a voltage of 200 kV. Representative sections were

9

photographed.

10 11

Insulin

content

and

encapsulation

efficiency.

The

insulin-loaded

12

nanoparticles were dispersed in 0.01 M HCl (10 mg in 10 ml) by magnetic stirring for

13

1 h at 1000 rpm and 25°C, and then centrifuged (Ultracentrifuge Optima LE-80K,

14

Beckman Coulter, USA) at 40,000 rpm for 7 h at 4°C.30 The supernatant was filtered

15

(polyvinylidene fluoride (PVDF) membrane filter, 0.45 µm; Durapore®, Millipore

16

Corporation, Ireland) and its insulin content was analyzed by HPLC technique. The

17

insulin content was expressed as a percentage of the weight of nanoparticles. The

18

insulin encapsulation efficiency was defined as the percentage of insulin encapsulated

19

in the nanoparticles with reference to the initial amount of insulin feed. At least

20

triplicates experiments were conducted and the results averaged.

21 22

Insulin release from nanoparticles. Nanoparticles (5 mg) were added to

23

tubes containing 5 ml of HCl/KCl buffer pH 1.2 or phosphate buffer pH 6.8, and the

24

dissolution experiments were conducted under sink conditions at 37 ± 0.2°C at 50

25

strokes/min in a shaker water bath (Memmert, Germany). At specified intervals of 30

26

min, 1 h and 2 h in simulated gastric medium (pH 1.2 buffer), and 4 h and 6 h in

27

simulated intestinal medium (pH 6.8 buffer), triplicate dissolution tubes were removed

28

and a 0.4 ml aliquot was sampled from each tube. Simulated intestinal samples were

29

acidified with 1 ml of 0.01 M HCl. All aliquots were filtered through 0.45 µm PVDF

30

membrane and assayed for insulin content by HPLC. Insulin release was calculated as

31

a percent of the total insulin load in the weighed amount of nanoparticles.

32 33

HPLC assay of insulin. Insulin was analysed by HPLC assay (Agilent 1200,

34

Agilent Technologies Inc., USA) at 25°C using an Agilent Zorbax SB-C18 reversed8 ACS Paragon Plus Environment

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1

phase column (4.6 × 250 mm; pore size 300 nm). The mobile phase consisted of (A)

2

0.03% trifluoroacetic acid in 90% H2O and 10% acetonitrile and (B) 0.03%

3

trifluoroacetic acid in 10% H2O and 90% acetonitrile. Analysis was conducted in

4

gradient mode using a mobile phase of 20:80 v/v A:B for 5 min, followed by run at

5

80:20 volume ratio of these mobile phase components for 10 min. The flow rate was

6

0.5 ml/min and the injection volume was 20 µL. Detection wavelength, stop time and

7

post time were set at 215 nm, 17.5 min and 3 min, respectively.

8

Differential scanning calorimetry (DSC). DSC thermograms (Pyris 6 DSC,

9

Perkin Elmer, USA) were obtained by crimping 3 mg of a solid sample in a standard

10

aluminium pan and heating the pan from 30 to 380°C at 10°C/min under a constant

11

purging of nitrogen at 40 ml/min. The characteristic peak temperature and enthalpy

12

values of endotherm and exotherm were recorded. At least triplicates were carried out

13

for each sample and the results averaged.

14 15

Bead size. Bead size was determined using a digital vernier caliper (Mitutoyo,

16

Japan). The length and breadth were measured of each bead and the mean of these two

17

dimensions calculated as the bead diameter. Ten beads were randomly selected for

18

measurement of the mean bead diameter of each batch.

19 20

Bead shape. The sphericity factor (SF) and aspect ratio (AR) were used to

21

estimate the roundness of the beads.31 The SF was calculated according to equation

22

(1):

23

SF = (dmax - dmin)/(dmax + dmin)

(1)

24

where dmax and dmin represented the maximum (length) and minimum (breadth)

25

diameters of the bead respectively. The SF value varied from 0 for a completely

26

symmetrical bead around its centre to approaching a unit value for an unshapely bead.

27

The AR was defined as the quotient of maximum diameter to minimum diameter of

28

beads as shown in equation (2):

29

AR = dmax/dmin

(2)

30

The AR value was 1 for a symmetrical bead and it increased as the bead became

31

elongated. Both SF and AR were characterized using a random sample of at least 15

32

beads for each batch prepared. 9 ACS Paragon Plus Environment


Page 10 of 43

1 2

Bead swelling, erosion and water uptake. Ten beads of each formulation had

3

their individual length and width measured using the digital caliper and their

4

respective weight characterized.32 Each bead was then placed in 5 ml of pH 1.2 buffer

5

to simulate gastric conditions and shaken at 50 rpm and 37°C for 2 h. The weight and

6

size of each wet bead were subsequently measured after removing its surface

7

moisture, achieved by rolling the bead gently over a dry petri dish till there was no

8

sign of moisture left in its immediate trial on the dish surface. The bead was then

9

oven-dried at 40ºC for 24 h and subsequently equilibrated to a constant weight by

10 11 12 13 14

storing in a desiccator at 25ºC. The swelling (SI), erosion (EI) and water uptake (WUI) indices of each bead were defined as: SI = (St- Si)/Si . 100%

(3)

where Si = initial dry bead diameter and St = wet bead diameter at time, t.

15 16 17

EI = Wi-Wt(d)/Wi . 100%

(4)

where Wi = initial dry bead weight and Wt(d) = dry weight of bead collected at t.

18 19 20

WUI = Wt-Wt(d)/Wt(d) . 100%

(5)

where Wt = wet weight of bead at t.

21 22

Insulin release from nanoparticles-loaded beads. The drug release profiles

23

of nanoparticles-loaded beads (25 mg of beads containing 5.52 mg nanoparticles with

24

0.1 mg insulin) transiting from acidic gastric milieu (2 h) to near-neutral intestinal

25

medium (4 h) were evaluated. This study utilized 3.5 ml of 0.1 M HCl to simulate

26

gastric fluid followed by its adjustment to pH 6.8 to simulate intestinal fluid through

27

adding 1.5 mL of 0.2 M of a tribasic sodium phosphate solution that had been

28

previously equilibrated to 37 ± 0.2°C. Aliquots (0.4 ml) were withdrawn at specified

29

intervals of 1 h and 2 h in simulated gastric medium, and 3 h and 6 h in simulated

30

intestinal medium. Simulated intestinal samples were acidified with 1 ml of 0.01 M

31

HCl. The aliquot was filtered through 0.45 µm PVDF membrane and had its insulin

32

content analyzed by HPLC. Fresh batches of beads were introduced into the test media

33

for sampling at each and every interval. The percentage of insulin release was


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1

calculated with respect to the total drug content in the beads. The drug content in the

2

beads was evaluated by incubating 100 mg of beads at 1000 rpm for 4 h at 25ºC in

3

phosphate buffer pH 6.8, followed by ultracentrifugation (Ultracentrifuge Optima LE-

4

80K, Beckman Coulter, USA) at 40,000 rpm for 1 h at 4ºC, and acidifying the

5

supernatant for HPLC analysis. Triplicates were conducted for each batch of beads

6

and the results averaged. When required, the aliquots obtained from the drug release

7

media were filtered using 0.45 µm PVDF membrane to remove the alginate bead gel

8

mass. The filtrates were subjected to particle size analysis to elucidate the status of

9

nanoparticle release from the alginate beads. In addition, beads were recovered at

10

dissolution intervals of 0 min, 120 min and 180 min, oven-dried at 40ºC for 48 h, and

11

analysed using FTIR and DSC techniques when necessary.

12 13

Cell Culture. HT-29 colon cancer cells purchased from the American Type

14

Culture Collection (ATCC, USA) were cultured in Eagle’s minimum essential

15

medium (MEM) (Sigma Aldrich, Germany), supplemented with 10 % fetal bovine

16

serum, 1.5 g/l sodium bicarbonate and 1% penicillin/streptomycin solution, at 37°C

17

under 5 % carbon dioxide and relative humidity of 95% (CO2CELL 170 incubator,

18

MMM, Germany).

19 20

Cytotoxicity of nanoparticles. The cytotoxicity of insulin-loaded ASAN and

21

insulin-loaded AC18N was evaluated in vitro with HT-29 cells using the MTT

22

assay.33-34 HT-29 cells (passage 6) were plated onto 96 well plates (Nunc, Denmark)

23

at a density of 5 × 104 cells/well and allowed to attach overnight in the incubator.

24

Following the removal of MEM, the cells were incubated with nanoparticle samples

25

(10 mg in 0.5 ml of phosphate buffer saline, pH 7.4 (PBS)) for 24 h. The cells were

26

then washed and incubated with 20 µl of 3-(4,5-dimethylthiazol-2-thiazolyl)-2,5-

27

diphenyl-2H-tetrazolium bromide (MTT; Sigma Aldrich, Germany; 5 mg/ml in PBS)

28

for 4 h in the absence of light. The supernatant was carefully discarded. The

29

intracellular blue formazan crystals were solubilized in 150 µl DMSO and quantified

30

at 570 nm by means of a microplate reader (Sunrise TECAN, Austria). The cell

31

viability was defined as absorbance ratio of treated sample to that of control (sample

32

obtained of cells incubated with MEM), expressed in percentage. The percent cell

33

death (CD) was calculated using the following equation:

34 11 ACS Paragon Plus Environment


Page 12 of 43

1

CD = [(absorbance of control-absorbance of treated cell line) / absorbance of control]

2

× 100

3

At least six replicates were conducted and the results averaged.

(6)

4 5

Cellular uptake. The cellular internalization profiles of FITC-labelled insulin-

6

loaded nanoparticles prepared with alginate-stearic acid and alginate-C18 conjugate

7

were visualized and quantified by laser scanning confocal electron microscope (Leica,

8

Germany). HT-29 cells (passage 6) were cultured at a density of 1 × 105

9

cells/fluoroDish and incubated until 70 to 80 % confluence. The culture medium was

10

discarded and the cells were washed twice with pre-warmed PBS before they were

11

incubated for 1 h at 37°C with FITC-labelled nanoparticles (1 ml, 2 mg/ml in PBS).

12

The test samples were aspirated and cells were washed twice with pre-warmed PBS

13

before cell imaging.

14

Laser scanning confocal electron microscope software (Advanced fluorescence

15

2.2.1 build 4842, Leica, Germany) was used to quantify the cellular uptake of

16

nanoparticles. Briefly, a cell of interest was selected using a polygon drawing tool. A

17

region next to the cell that had no fluorescence was similarly selected as the

18

background. The same process was repeated in the field of view. At least 6 images

19

from three different experiments were analyzed. Within each image, 5 to 10 random

20

regions of interest were selected. The total cell fluorescence (TCF) was calculated

21

using the following equation:

22 23

TCF=integrated density-(area of selected cell × mean fluorescence of background) (7)

24 25 26

In the same study, the endocytic pathways of nanoparticles were examined

27

through introducing and incubating the HT-29 cells (5000 cells/fluoro-dish with 80 %

28

confluence) with 1 ml of pharmacological membrane entry inhibitors. Seven µg/ml

29

chlorpromazine (Calbiochem, USA) were used to inhibit the formation of clathrin

30

vesicles,35-36 200 µM genistein (Calbiochem, USA) to inhibit caveolae pinching,36-37

31

500 nM Wortmannin (Calbiochem, USA) to inhibit phosphatidylinositol 3-kinase

32

(macropinocytosis)35,38 and 25 µg/ml nystatin (Calbiochem, USA) to interact with

33

cholesterol (lipid rafts)39 for 1 h at 37oC. The inhibitor solutions were then removed 12 ACS Paragon Plus Environment

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1

and the PBS pH 7.4 suspension of nanoparticles (1 mg/ml) was added and further

2

incubated for 1 h at 37oC. The cells were then washed twice with pre-warmed PBS pH

3

7.4 and subjected to confocal microscopy analysis.

4 5

In vivo characterization. Healthy male Sprague dawley rats (Lafam, UiTM,

6

Malaysia), aged 3 months and weighing 200 to 250 g, were acclimatized for 7 days in

7

individual housing under 12 h light/dark cycle with deionized water and standard

8

pelletized food (Gold Coin Enterprise, Malaysia) given ad libitum. The ambient

9

temperature was set at 25 ± 2°C with relative humidity maintained at 55 ± 5 %. All

10

experiments were conducted in accordance to the university ethics regulations

11

adapting the international guidelines (OECD Environment, Health and Safety) on the

12

conduct of animal experimentation.

13 14

Mucus penetration. FITC-labelled nanoparticles were prepared by covalently

15

conjugating the FITC to the alginate backbone following by nanospray drying of a

16

solution of the FITC-alginate. Rats (n = 3) were fasted for 12 h prior to sacrificed by

17

cervical dislocation technique. The duodenum was isolated and excised longitudinally

18

to obtain a tissue segment of 1 cm in length and width. The tissue segment was placed

19

in a glass petri dish with the luminal surfaces of epithelium facing upwards. An

20

accurately weighed 0.2 mg sample of FITC-labelled nanoparticles was dispersed in 20

21

µL water and transferred onto the tissue epithelium by means of an electronic

22

micropipette to cover an area of approximately 1 cm2. After 30 min, the tissue

23

segment was sectioned at 90° to the epithelial surface using a cryostat (CM1850 UV-

24

1-1, Leica, Germany). The tissue segment was sealed to prevent moisture losses by

25

placing a cover slide over the sample. Images of epithelium level were captured using

26

a confocal electron microscope (Leica, Germany). The mucus layers were identified

27

as a zone less dense than that of adventitia with rich blood supplies. The fluorescence

28

intensity of penetrating nanoparticles embedded in the mucus layer of epithelium was

29

computed by laser scanning confocal electron microscope software (Advanced

30

fluorescence 2.2.1 build 4842, Leica, Germany).

31 32

Blood glucose and insulin. Diabetes was induced in the rats by a single

33

intraperitoneal injection of 60 mg/kg streptozotocin in isotonic saline solution. Two

34

weeks following the streptozotocin administration, rats with fasted blood glucose 13 ACS Paragon Plus Environment


1

levels in the range of 13.89 to 16.67 mmol/L were used for experiments. The diabetic

2

rats were randomly divided into seven groups (n = 6/group). Control groups consisted

3

of rats given 1 ml of saline solution orally (negative control), blank AC18N orally,

4

blank AC18N embedded in tripolyphosphate-crosslinked chitosan-oleic acid

5

conjugate-coated calcium alginate beads (CCAB) orally, or 0.036 mg/kg (equivalent

6

to 1 IU/kg) of insulin solution by subcutaneous injection (positive control; 0.3 ml).

7

Treatment groups comprised of rats administered with 0.54 mg/kg (equivalent to 15

8

IU/kg) of insulin in the form of oral insulin solution (1 ml), insulin-loaded AC18N

9

orally, or insulin-loaded AC18N embedded in CCAB orally. Nanoparticles and beads

10

were administered in the form of a hard gelatin capsule with reduced volume (length =

11

8.16 mm instead of 22.94 mm) for ease of swallowing.40 All rats were fasted for 12 h

12

prior experiments.

13

The rats were anesthetized by means of ketamine-xylazine intraperitoneal

14

injection (15 mg ketamine/200g rat and 2 mg xylazine/200 g rat) for the collection of

15

blood samples (200 µl) from the retroorbital plexus at specified time intervals. The

16

blood glucose level was determined using a glucometer (Ascensia Elite, Bayer

17

Corporation, Belgium), and the changes in blood glucose level were expressed as a

18

percent relative to the baseline blood glucose concentration at 0 h. To determine the

19

plasma insulin concentration, the 24-h blood samples were collected and allowed to

20

clot with sera obtained by centrifugation of the samples at 4100 rpm for 20 min at

21

25°C. Serum samples were preserved at -20°C until further analysis using the insulin

22

enzyme immunoassay kit (A05105-96 WELLS; SPI-BIO, France). Insulin

23

concentrations were measured at 405 nm using a plate reader (Gen 5 microplate

24

reader, BioTek, USA).

25 26

Statistics. The results are expressed as mean and standard deviation. Pearson

27

correlation, Student’s t-test and analysis of variance (ANOVA)/post hoc analysis by

28

Tukey HSD were carried out using SPSS software 16.0. Statistical significance was

29

denoted by p < 0.05 unless otherwise stated.

30 31

Results and discussion

32 33

Synthesis of alginate-C18 conjugate. The alginate-C18 conjugate was

34

synthesised with the aim to introduce a hydrophobic hydrocarbon segment to the 14 ACS Paragon Plus Environment

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1

hydrophilic polymeric backbone. Figure 1 shows the schematic diagram of

2

conjugation reaction, and the FTIR and NMR spectra of alginate, 1-bromooctadecane

3

and alginate-C18 conjugate. Successful conjugation of alginate with C18 was

4

indicated by the appearance of new FTIR bands at 1739.8 ± 1.8 cm-1, ascribing to the

5

formation of ester bond between alginate and 1-bromooctadecane, and 2923.3 ± 0.4

6

cm-1 and 2854.8 ± 1.8 cm-1 attributable to the availability of the C18 hydrocarbon

7

chain. The FTIR spectrum of alginate was characterized by bands at 1417.7 ± 0.5 cm-1

8

and 1613.4 ± 1.0 cm-1 ascribing to C=O moiety,41 and 3442.3 ± 2.4 cm-1 in association

9

with its O-H functional group.42 The conjugation of alginate with C18 led to reduced

10

wavenumber values of these peaks possibly due to an increase in intramolecular

11

interaction propensity between the adjacent C=O and O-H moieties.43

12

NMR analysis of alginate, 1-bromooctadecane and alginate-C18 conjugate

13

showed that the alginate was characterized by chemical shifts at 4.9, 4.3 and 4 ppm

14

attributable to the H1, H5 and H3 moieties of guluronic acid, respectively (Figure 1).44

15

The formation of alginate-C18 conjugate was indicated by appearance of peaks in the

16

chemical shift range of 0.8-2.6 ppm attributable to the CH3 and CH2 groups in the C18

17

chain.45

18



1 2

Figure 1. Schematic diagram of alginate-C18 conjugation reaction, and FTIR and

3

NMR profiles of (a) alginate, (b) 1-bromooctadecane and (c) alginate-C18 conjugate.

4 5

Preparation of insulin-loaded nanoparticles. Insulin-loaded nanoparticles

6

prepared using alginate, alginate-stearic acid, and alginate-C18 conjugate were

7

fabricated by nanospray drying technology. These nanoparticles were all spherical in

8

shape (Table 1), but the ASAN were larger in size than the AN and AC18N (Table 1).

9

The incorporation of the low bulk density stearic acid (0.84 g/ml)46 could have

10

expanded the size of the nanoparticles. Conversely, the conjugation of C18 did not

11

affect the size of the alginate nanoparticles. The insulin encapsulation efficiency and

12

insulin content of ASAN were higher than those of AN and AC18N (Table 1;

13

ANOVA: p < 0.05). It was conferred by deterred insulin leaching due to hydrophobic

14

effect of stearic acid that was freely dispersed instead of confined to specific segments


Page 16 of 43

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1

of alginate chains in these nanoparticles.

2

Nanoparticles of the polyanionic alginate were, as expected, negatively

3

charged due to the availability of free surface COO- moieties (Table 1). Processing of

4

alginate through physical blending with stearic acid or conjugating with C18 followed

5

by nanospray drying translated to the formation of nanoparticles with reduced zeta

6

potentials, particularly for the AC18N (Student’s-t-test: AN vs ASAN: p = 0.02; AN

7

vs AC18N: p = 0.00; Table 1). C18 conjugation occurred at the COO- sites in the

8

alginate, and the depletion of COO- groups was reflected in the significantly lower

9

zeta potential values of the AC18N. These nanoparticles were envisaged to experience

10

the lowest level of inter-molecular electrostatic repulsion between the polymer chains,

11

which would further explain why these nanoparticles were smaller in size compared

12

with the ASAN, despite the incorporation of the C18 hydrocarbon chains into both

13

types of nanoparticles.


Molecular Pharmaceutics 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47

Page 18 of 43

Table 1 Formulation and physicochemical characteristics of insulin-loaded AN, ASAN and AC18N. Formulation

Physicochemical characteristics

Nanoparticle

Alginate

Stearic

Alginate-

Insulin

Particle size

type

(mg)

acid

C18

(mg)

(nm)

(mg)

conjugate

PDI

Zeta

Drug

Drug

TEM image

potential

content

encapsulation

(magnification ˃ 1000 ×)

(mV)

(%)

efficiency

(mg)

Insulin-loaded

100

___

___

(%)

5

AN

Insulin-loaded

100

10

___

5

ASAN

Insulin-loaded AC18N

___

___

100

5

513.00 ±

0.54 ±

-45.17 ±

3.70

44.38 ±

19.00

0.02

2.38

±0.15

1.40

618.87 ±

0.33 ±

-40.13

6.44 ±

76.69 ±

6.57

0.04

±1.56

0.92

10.92

522.50 ±

0.74 ±

-35.67 ±

3.77 ±

44.87 ±

66.47

0.18

0.70

0.13

1.55

18

ACS Paragon Plus Environment

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1

Drug release profiles from nanoparticles.

In vitro drug release study

2

showed undetectable levels of insulin release from the AN even after 6 h of incubation

3

in the dissolution medium. The alginate has a pKa value of 3.65,47 while the

4

amphoteric insulin molecule has an isoelectric point of 5.3.48 The nanoparticles were

5

prepared by mixing the negatively charged alginate (dissolved in NaOH) with the

6

positively charged insulin (dissolved in HCl). Complex formation with the alginate

7

would prevent the insulin from being released from the formed matrix.49 This was

8

supported by DSC analysis of alginate, insulin and insulin-loaded AN. The

9

nanoparticles were characterized by an endotherm of lower peak temperature and

10

higher melting enthalpy (419.1±107.7 J/g) than those of alginate (132.2 ± 7.0 J/g) and

11

insulin (136.2 ± 41.4 J/g) (Figure 2). The introduction of insulin into the alginate

12

matrix could mutually reduce the strength of physicochemical interaction of the

13

individual chemical species. It conferred a higher extent of inter-species interaction

14

between the alginate and the insulin through dispersion of insulin in the alginate

15

matrix. FTIR spectrum of the insulin-loaded AN also exhibited a shift to reduced

16

wavenumber at 3081.3 ± 11.9 cm-1 which might be attributable to alginate-insulin

17

interaction via O-H and/or N-H moieties (Figure 2). Based on its in vitro insulin

18

release profile, the insulin-loaded AN were not subjected to further evaluation in the

19

cell-based and in vivo experiments.

20

Addition of stearic acid to the alginate nanoparticles resulted in a prompt

21

release of insulin within the first hour of dissolution experiment (Figure 3).

22

Prolongation of the dissolution time to 2 h tend to lead to a decline in fraction of

23

insulin released (Student’s t-test: 1 h vs 2 h: p = 0.07). A similar observation was

24

noted when the drug release study was conducted in the simulated intestinal medium

25

(Figure 3; Student’s t-test: 4 h vs 6 h: p = 0.01). The stearic acid was available as free

26

molecules in the nanoparticles, as evidenced by the existence of an endotherm

27

ascribable to the fatty acid at 54.9 ± 0.1ºC in the DSC thermogram of the

28

nanoparticles (Figure 2). Similarly, FTIR spectrum for the nanoparticles showed

29

transmission bands at 2919.8 ± 0.2 cm-1 and 2851.4 ± 0.3 cm-1 that were characteristic

30

of the C-H moiety of stearic acid (Figure 2). The dispersion of the hydrophobic stearic

31

acid in the solid hydrophilic alginate matrix could induce immiscible zone formation

32

as a result of physical incompatibility between the two domains.50 The stearic acid

33

also appeared to disrupt the polymer-polymer, drug-drug and polymer-drug

34

interactions in the nanoparticles. DSC thermogram of the insulin-loaded ASAN 19 ACS Paragon Plus Environment


Page 20 of 43

1

showed a broad endotherm with lower peak temperature at 121.7 ± 6.7ºC compared

2

with the corresponding endotherm of the insulin-loaded AN (Figure 2). The

3

wavenumbers of FTIR peaks corresponding to O-H/N-H moiety (3436.7 ± 2.9 cm-1)

4

and C=O moiety (1621.3 ± 0.9 cm-1) for the ASAN were higher than those for the AN

5

(Figure 2). The combined effects of immiscible zone formation and disruption of

6

alginate-insulin interactions would promote insulin release from the ASAN. With

7

time, the released insulin might be re-adsorbed onto the alginate domain of the

8

nanoparticles via interaction between the O-H, C=O and/or N-H moieties of the

9

polymer and drug. The availability of free insulin in the dissolution media was thus

10

decreased.

11

The drug release propensity of the AC18N, despite their smaller size and

12

therefore larger specific surface area available for drug dissolution, was lower than

13

that of ASAN (Figure 3; Table 1). Pearson correlation analysis of drug release profiles

14

of AN, ASAN and AC18N indicated that both drug content and drug encapsulation

15

efficiency could have partially accounted for the said observation (r = 0.63 – 0.86).

16

The conjugation of alginate with C18 could bring about the formation of a relatively

17

compact nanoparticulate matrix structure possibly due to amphiphilic character

18

enhancement than nanoparticles prepared with alginate physically mixed with stearic

19

acid. The level of interaction between the alginate-C18 conjugate and insulin via O-

20

H/N-H and C=O moieties was higher than that between alginate-stearic acid and

21

insulin. This was indicated by FTIR analysis where the AC18N were characterized by

22

peaks with lower wavenumbers at 3429.2 ± 8.4 cm-1 (O-H/N-H) and 1615.5 ± 0.4 cm-1

23

(C=O) (Figure 2). These nanoparticles also exhibited

24

temperatures, at 150.2 ± 4.8ºC and 154.1 ± 5.7ºC, than the ASAN (Figure 2). Further,

25

the conjugated hydrocarbon chain was expected to undergo a lower degree of leaching

26

than the freely dispersed stearic acid within the nanoparticles. It was more readily

27

available to act as a hydrophobic domain in retarding drug dissolution thereby

28

reducing drug release from the matrix. The released insulin did not appear to be re-

29

adsorbed onto the AC18N matrix with prolonged dissolution in simulated intestinal

30

medium at pH 6.8 (Figure 3). Attachment of the C18 meant the AC18N had a low

31

quantum of COO- moiety, which could have partially prevented the dissolved insulin

32

from readsorption onto the alginate nanoparticles.

33 34 20 ACS Paragon Plus Environment

higher melting peak

Page 21 of 43 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60


1 2 3 4 5 6 7 8 9 10 11 12 13 14 21 ACS Paragon Plus Environment


1 2

Figure 2. DSC and FTIR profiles of (a) insulin, (b) alginate, (c) stearic acid, (d)

3

insulin-loaded AN, (e) insulin-loaded ASAN and (f) insulin-loaded AC18N. T/°C,

4

∆H/J/g.

5 6


Page 22 of 43

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1 2 3

Figure 3. Drug release profiles of insulin-loaded (a) ASAN and (b) AC18N in

4

simulated gastric medium pH 1.2 and simulated intestinal medium pH 6.8.

5 6

Cytotoxicity, mucus penetration and intracellular trafficking. In vitro

7

cytotoxicity, as well as the capacity for mucus penetration and intracellular trafficking

8

of the ASAN and AC18N were evaluated. These nanoparticles were envisaged to act

9

as carriers to transfer the insulin load transmucosally into the systemic circulation.

10

Their ability to penetrate the intestinal mucus to reach the intestinal epithelium and 23 ACS Paragon Plus Environment


1

undergo transepithelial uptake are therefore important factors influencing their

2

efficacy in vivo.

3

MTT assays indicated that both the insulin-loaded ASAN and insulin-loaded

4

AC18N were characterized by low levels of cytotoxicity (Figure 4a). The viability of

5

HT-29 cells was not affected by the nanoparticles (ANOVA, p > 0.05). The mucus

6

penetration capacity of AC18N appeared to be higher (Student’s-t-test: p = 0.06) than

7

that of ASAN (Figure 4b). The AC18N were characterized by a lower magnitude of

8

negative charges at the particulate surfaces. Being smaller in size and having a lower

9

level of negative surface charges (Student’s-t-test: size, p < 0.05; zeta potential, p

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