Journal of Controlled Release 186 (2014) 11–21

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Journal of Controlled Release

j ourna l homepage: www.e lsev ie r .com/ locate / jconre l

Review

Particle designs for the stabilization and controlled-delivery of proteindrugs by biopolymers: A case study on insulin

Hui-Peng Lim a, Beng-Ti Tey a,b, Eng-Seng Chan a,b,⁎a Chemical Engineering Discipline, School of Engineering, Monash University, Jalan Lagoon Selatan, Bandar Sunway 46150, Selangor, Malaysiab Multidisciplinary Platform of Advanced Engineering, Monash University, Jalan Lagoon Selatan, Bandar Sunway 46150, Selangor, Malaysia

⁎ Corresponding author at: Chemical Engineering Di5514 5821; fax: +60 3 5514 6207.

E-mail addresses: lim.hui.peng@monash.edu (H.-P.

http://dx.doi.org/10.1016/j.jconrel.2014.04.0420168-3659/© 2014 Elsevier B.V. All rights reserved.

a b s t r a c t

a r t i c l e i n f o

Article history:Received 16 January 2014Accepted 23 April 2014Available online 6 May 2014

Keywords:Protein drugInsulinEncapsulationBiopolymerOral delivery

Natural biopolymers have attracted considerable interest for the development of delivery systems for proteindrugs owing to their biocompatibility, non-toxicity, renewability and mild processing conditions. This paperoffers an overview of the current status and future perspectives of particle designs using biopolymers for thestabilization and controlled-delivery of a model protein drug — insulin. We first describe the design criteria forpolymeric encapsulation and subsequently classify the basic principles of particle fabrication as well as theexisting particle designs for oral insulin encapsulation. The performances of these existing particle designs interms of insulin stability and in vitro release behavior in acidic and alkalinemedia, as well as their in vivo perfor-mance are compared and reviewed. This review forms the basis for future works on the optimization of particledesign and material formulation for the development of an improved oral delivery system for protein drugs.

© 2014 Elsevier B.V. All rights reserved.

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 122. Design criteria for polymeric encapsulation of oral insulin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12

2.1. Formulation and processing conditions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 122.2. Particle size . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 132.3. Protection against gastric solution in the stomach . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 132.4. Controlled-release at the targeted site . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13

3. Natural biopolymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 134. Basic principle for forming hydrogel particles and loading of insulin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13

4.1. Liquid–air methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 134.2. Liquid–liquid methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 144.3. Comparison between dispersion methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 144.4. Insulin loading method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 14

5. Existing particle designs and their performance . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 145.1. Beads with a single polymer . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 165.2. Beads with blended polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 175.3. Coated beads . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 175.4. Coated beads with emulsified insulin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 185.5. Colloidosomes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 185.6. Hydrogel-coated particles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 185.7. Polyelectrolyte complexes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 185.8. Polymer-insulin complexes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 185.9. Effect of particle size . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 19

scipline, School of Engineering, Monash University, Jalan Lagoon Selatan, Bandar Sunway 46150, Selangor, Malaysia. Tel.: +60 3

Lim), tey.beng.ti@monash.edu (B.-T. Tey), chan.eng.seng@monash.edu (E.-S. Chan).

12 H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

6. Concluding remarks and future perspectives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 19Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 20References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 20

1. Introduction

Intensive research investigating the therapeutic potential of proteinsand peptides has introduced a substantial number of protein drugs tothe pharmaceutical industry. For instance, various health-jeopardizingdiseases, such as cancer, infectious diseases, and autoimmune diseases,can now be treated using protein drugs, including aldesleukin, humangrowth hormone, bone morphogenic protein, and others [1]. One ofthe most prominent diseases treated with protein drugs is diabetesmellitus. Diabetes is a hyperglycemia-causing disease that currentlyaffects approximately 382 million people worldwide [2]. Patientsoften have to rely on exogenous insulin to regulate their bloodglucose levels for a healthier life.

Insulin was first discovered in the canine pancreas by Banting andBest in 1921 [3]. Throughout the years, insulin has been successfullyisolated from other sources, such as porcine and bovine pancreases.Nevertheless, insulin from animal sources often caused allergic reac-tions in humans. Owing to the advancement of biotechnology, humaninsulin can now be produced in mass quantities through recombinantDNA technology [4]. Today, a variety of insulin formulations [5] withdifferent onsets and durations of action are available on the market tomeet the different needs of diabetes patients.

Regardless of the formulation type, exogenous insulin is commonlyadministered via subcutaneous injection into the fatty areas of thelower abdomen, thigh, buttocks, and upper arm [6]. A typical diabeticpatient requires more than 60,000 injections throughout his or her life[7]. This route of administration, however, has several drawbacks,such as the inconvenience of multiple injections, occasional hypoglyce-mia due to insulin overdose, a risk of infection at the injection sites, andmost importantly, poor patient compliance with injections [4,8,9].

Other non-invasive means of insulin delivery, including ocular,vaginal, rectal, buccal or sublingual, oral, and nasal are currently beinginvestigated. Most of the proposed means, however, involve the use ofabsorption enhancers, such as bile salts and surfactants, to overcomethe thick mucosal layers present in the tissues, which may causecomplications if an overdose is administered [10–12]. For this reason,oral administration, which has a high level of patient acceptance andprovides a larger area for absorption is a highly desirable alternative.Upon oral administration, a high level of insulin is directly channeledfrom the intestine to the liver and then to the portal blood, mimickingthe endogenous insulin secretion pattern of the β-cells of the pancreas[13].

Nevertheless, the oral delivery of insulin involves uptake of the drugin the gastro-intestinal tract (GIT), which has abundant proteolyticenzymes and varying pH conditions. For instance, the pH in the stomachcan vary from 1 to 5, whereas the pH in the intestine or colon can varyfrom 6.4 to 7.5 [14]. Alternative approaches to deliver insulin effectivelyin the GIT are being investigated. For example, chemical modification ofinsulin molecules [15], co-administration of insulin with absorptionenhancers [16,17] or protease inhibitors [18,19], and association of insulinwith a drug delivery system, such as polymeric micro and nano carriers[20–23], and lipid-based carriers such as liposomes [24–27] and solidlipid nanoparticles (SLN) [28], have been proposed. Insulin-loadedliposomes and SLN were reported to exhibit a hypoglycemic effectwhen orally administered to diabetic rats [24,25]. Nonetheless, theefficacy of the liposomes and SLN was often reduced due to thedegradation of the lipid-based carriers in the GIT [24,25].

Polymeric encapsulation usinghydrophobic or hydrophilic polymershas received considerable attention [5,29,30]. Researchers often resort

to synthetic hydrophobic polymers because of their tunable degrada-tion properties. Examples of synthetic hydrophobic polymer frequentlyused in insulin oral delivery systems are those consisting of α-hydroxyacids, such as polyglycolic acid (PLA), polylactic acid (PGA), andpoly(lactic-co-glycolic acid) (PLGA) [31–34]. The degradation rate ofPLGA, for example, can be varied by altering the co-polymer ratio.Nevertheless, the degradation of the polymer causes an acidification ofthemicroenvironment in themicrospheres [35–39], resulting in insulindegradation similar to that observed upon incubation in buffer solutionat pH 2 [35].

Pharmaceutical companies around the globe are developing oralinsulin products that are mainly in capsule or tablet form. Most of theoral insulin products have adjuvants that protect the insulin frompremature release in the stomach and absorption enhancers that pro-mote intestinal uptake. A few products are currently in the latter stageof clinical trials [40]. In particular, efficacy of the enteric-coatedinsulin-loaded capsule containing adjuvants (ORMD-0801) developedby Oramed Pharmaceuticals has been shown in phase II clinical trialson the basis of suppressed rise of glucose level in fasting subjects [41,42]. Capsulin, which was developed by Diabetology Ltd., is anotherproduct that has undergone testing in phase II clinical trials, and itwas reported to be capable of controlling blood glucose levels for anextended time period [43].

In recent years, natural-source polymers (i.e., biopolymers),which aregenerally hydrophilic, have received much attention from researchersdue to their low cost, renewability, biodegradability, biocompatibility,and non-toxicity that makes them safe for human consumption. Soniaand Sharma [13] previously reviewed the suitability of different types ofbiopolymers as polymeric material for oral insulin delivery systems. Inthis review, we compare the stabilization and delivery performances ofbiopolymeric oral insulin delivery systems by outlining the variousparticle designs and fabrication techniques reported in the literature.The extent of success of the various particle designs in stabilizing anddelivering insulin was compared in terms of the insulin stability andin vitro release behavior in acidic and alkaline media, as well as theirin vivo performance.

2. Design criteria for polymeric encapsulation of oral insulin

Encapsulation is the process of confining active compounds within amatrix in particulate form to achieve desirable effects [44]. In thecontext of oral insulin delivery, the goal of encapsulation is to provideprotection and controlled-delivery to the encapsulated insulin in theGIT. Prior to designing the polymeric oral insulin delivery systems,several criteria that could affect the efficacy of insulin delivery, especiallyfor maintaining the stability and increasing bioavailability, must be givenproper consideration.

2.1. Formulation and processing conditions

A good delivery system should contain biologically active insulinupon the encapsulation process. The native structure of insulin is easilydisrupted upon exposure to extreme acidic pH (i.e., pH b 3) [45–47],high temperature (i.e., 50 to 80 °C) [48], and organic solvents (i.e.,methylene chloride) [49]. The disruption of either level of the hierar-chical structure of insulin can result in insulin degradation, ultimatelyleading to the loss of its blood glucose lowering ability [4,29,50]. There-fore, mild formulation (i.e., near neutral pH and zero-usage of organic

13H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

solvents) and processing conditions (i.e., room temperature) aregenerally preferred.

2.2. Particle size

Depending on the size of the polymeric carrier, insulin can be deliv-ered through different mechanisms in the GIT. For instance, particlessmaller than 2 μm can directly diffuse through the intercellular spacesof the intestinal epithelial cells or are absorbed through Peyer's patches,which are abundant in the distal intestine [50]. Once the particles are inthe bloodstreamand traveling to other organs in the body, the release ofinsulin can occur through either diffusion or polymer degradation. Forparticles larger than 2 μm, the release of insulin can occur in the intestinallumen, before being absorbed into the bloodstream through themicrovillithat are abundant on the intestinal wall.

2.3. Protection against gastric solution in the stomach

Ideally, an oral insulin delivery system should be able to protect theencapsulated insulin from theharsh acidic conditions in the stomach. Toavoid hydrolysis and proteolytic degradation of the encapsulated insulinmolecules, the inward permeation of hydrogen ions and proteolyticenzymes from the gastric solution should be minimized, if not totallyprevented. Zero or minimal release of the encapsulated insulin in thestomach is also desired to ensure a higher release of biologically activeinsulin at the targeted site of absorption later on.

2.4. Controlled-release at the targeted site

An ideal insulin delivery system should control the release of theencapsulated insulin at the targeted site of absorption. The release ofbiologically active insulin in the ileum of the small intestine is preferredbecause the ileum has a better capacity for insulin absorption comparedto the duodenum and jejunum [21,51]. Hence, gradual and sustainedrelease of biologically active insulin in the ileum is desired to ensurebetter control over the patient's glycemic level. However, a sudden orburst release of insulin could lead to hypoglycemia in a diabetic patient.

3. Natural biopolymers

Biopolymers from animal- or plant-based derivatives are generallyhydrophilic in nature. Their gel-forming property is often exploited toentrap insulin in an insoluble semi-permeable matrix called a hydrogel.The hydrogels formed can be molded into various shapes, such ascubical, cylindrical, and slab, although hydrogels in particulate form,which are generally referred to as beads, capsules or particles (see Fig. 1for example), are more commonly used for oral insulin encapsulation.

Biopolymers that have been frequently investigated as deliverysystems for oral insulin include alginate, chitosan, pectin, gelatin, anddextran sulfate [13]. Most biopolymers are safe for human consumptionand have been listed in the Code of Federal Regulations (21 CFR Parts

Fig. 1. SEM images of insulin-loaded beads or particles. (a) Alginate microbeads and(b) chitosan nanoparticles. Reprinted from refs. [52],©2006, withpermission fromElsevier; and [53],©2002, withpermission from Wiley-Liss, respectively.

184) as generally regarded as safe (GRAS) by the Food andDrug Admin-istration (FDA). They also possess a mucoadhesive property that isdesirable for improving the bioavailability of the insulin. Of all the bio-polymers, alginate and chitosan are widely studied due to their mildgelling conditions [54].

Briefly, alginate is a polysaccharide derived from marine plants thatis composed of (1→ 4)-linked β-D-mannuronic (M) and α-L-guluronicacid (G) residues and is polyanionic in nature. Alginate ismucoadhesivebecause of the presence of carboxyl functional groups in the polymericchain. The presence of such strong hydrogen-bonding groups in thepolymeric chains enhances interactions with the hydrophobic mucosallayer of the intestinal wall [55]. This enhanced interaction in turnprolongs the residence time of the polymeric vesicles at the intestinalwall, and thus results in increased absorption efficiency. Alginate canform a hydrogel at room temperature under mild condition throughtwo mechanisms: firstly, ionotropic gelation in the presence of cross-linkers such as Ca2+ and secondly, polyelectrolyte complexation withan oppositely charged biopolymer such as chitosan.

Chitosan, on the other hand, is a polysaccharide copolymer of β(1 → 4) linked glucosamine and N-acetyl glucosamine, that is derivedfrom the chitin of crustacean shells through a partial deacetylationprocess [56,57]. The polycationic chitosan is able to adhere to the anionicmucosal layer of the intestinal wall and trigger the opening of tightjunctions between the epithelial cells to enhance permeation [58,59].Chitosan can form a hydrogel structure through ionotropic gelation inthe presence of cross-linkers, such as tripolyphosphate (TPP) andpoly(γ-glutamic acid) (γ-PGA), or through covalent cross-linking in thepresence of glutaraldehyde. In addition, chitosan can form an insolublematrix through polyelectrolyte complexation with polyanions such asalginate or by precipitation at a pHabove its pKa value (i.e., about pH 6.5).

Other less commonly used biopolymers such as pectin can also formhydrogels through ionotropic gelation in the presence of cations. Gelatin,on the other hand, forms hydrogels through thermal gelation andcovalent cross-linking. The thermally responsive polymer chains ofgelatin coil into helices to form a hydrogel structure upon coolingbelow room temperature (i.e., 25 °C) [60]. The gelatin hydrogel structurecan be strengthened through covalent cross-linking in the presence ofglutaraldehyde. However, the use of glutaraldehyde in drug deliverysystems should be avoided, if possible, due to its toxicity. The differentgelation mechanisms for the aforementioned biopolymers are summa-rized in Table 1.

4. Basic principle for forming hydrogel particles and loading ofinsulin

In general, hydrogel encapsulation techniques involve two processingsteps, namely the dispersion of biopolymer solution into liquid dropletsfollowed by a solidification (or gelation) process (see Fig. 2). Droplet dis-persion techniques can be categorized into liquid–air and liquid–liquidmethods. The size of the particles formed is determined by the dispersionmethod used (see Fig. 3). While there are many known dispersionmethods, this review, only highlights the dispersion methods that havebeen used for the encapsulation of oral insulin by biopolymers.

4.1. Liquid–air methods

In a typical liquid–air method, biopolymer solution is dispersed intodroplets in the air phase using a nozzle-based system. The dropletsformed would fall under the force of gravity to a liquid bath placedunderneath the nozzle, and the gelation of the droplets takes placein this liquid bath. The gelation method employed depends on thebiopolymer used, as described in the earlier section.

Liquid–air methods normally produce large beads with a mean sizeranging frommicrometers to millimeters (see Fig. 3). Extrusion drippingis the simplest and most commonly used technique in this category. Inthis method, biopolymer solution is extruded drop-wise under the

Table 1Gelation methods of various biopolymers used for oral insulin encapsulation.

Biopolymer Ionotropic gelation Polyelectrolyte complexation Precipitation Covalent cross-linking Cooling

Alginate In the presence of calcium ions With chitosan N/A N/A N/AChitosan In the presence of TPP/γ-PGA With alginate or dextran sulfate At pH ≥ 6.5 In the presence of glutaraldehyde N/AGelatin N/A N/A N/A In the presence of glutaraldehyde At temperature b25 °CPectin In the presence of calcium ions N/R N/A N/A N/A

Notes: N/A = Not applicable; N/R = Possible but not reported.

14 H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

influence of gravitational force to produce beads of millimeter size withnarrow size distribution [61,62]. Beads in the sub-millimeter range canbe produced by promoting the droplet breakup at the dripping tip usingan external force, such as vibration. The air atomization technique hasalso been used to produce microbeads [63,64], but the size distributionof the particles formed is wider than that produced using the extrusionmethod.

Another technique in the liquid–air category that is noteworthyinvolves the dispersion of liquid droplets containing a very low concen-tration of cross-linking agents (i.e., 0.1 to 0.2%w/v in the case of ionotropicgelation) or oppositely charged polyelectrolyte (i.e., 0.05 to 0.09% w/v inthe case of polyelectrolyte complexation) into a solution containing avery low concentration of biopolymer (i.e., generally b 0.08% w/v). Thistechnique, which is normally carried out using simple agitation, is usedto produce nano-sized particles [65–68].

4.2. Liquid–liquid methods

In contrast to liquid–air methods, liquid–liquid methods involve theformation of liquid droplets by dispersing the biopolymer solution in animmiscible continuous liquid phase. This method, which is also knownas emulsification, results in the formation of a water-in-oil (W/O)emulsion. Typical “oil phases” used are vegetable oils, paraffin oil ororganic solvents such as chloroform. The droplets can subsequently besolidified by initiating the gelation process. In this method, the choiceof the emulsification technique and the surfactant formulation play acritical role in determining the size and size distribution of the particlesproduced. For instance, emulsification by mechanical agitation canproduce particles over a large range of mean sizes, but with broad sizedistributions [69]. Membrane emulsification can produce particles ofnarrower size distributions [70,71], whereas high pressure homogeniza-tion can produce particles in the nano-size range [72].

4.3. Comparison between dispersion methods

Both dispersion methods mentioned above have their own advan-tages and drawbacks. The particle sizes and size distributions are the

Fig. 2. Categorization of encapsulation techniques (i.e., droplet dispersion and gelation) forbiopolymeric oral insulin delivery systems.

main determining factors for the selection of a dispersion method.Several other factors, such as the process economics and scalabilitymust also be considered, especially for industrialization purposes. Inliquid–air methods, the polymer droplets that are formed in a “drop-by-drop” basis can be scaled-up usingmultiple nozzles. To ensure the forma-tion of uniform droplets at each nozzle, the nozzles can be equippedwitha control system to precisely control the liquid flow rate and dropletgeneration behavior. However, this sophisticated setup is associatedwith a high capital cost for production. Liquid–liquid systems, on theother hand, only require a simple stirred tank for high volumeproduction.Nevertheless, the emulsificationmethodoften involves the use of paraffinoil, vegetable oil, chloroform, or toluene as the immiscible phase. Withthe exception of vegetable oil, the use of the organic solvents, as well asthe large amounts of oil-solubilizing chemicals used during the washingsteps increases operating costs and also leads to concerns regarding thesafety of the produced particles. Furthermore, the repetitive washingand rinsing steps required to remove the excessive surface oil prolongthe processing time and generate large amounts of waste oil.

4.4. Insulin loading method

Insulin can be loaded into particles through two methods, namelypre-gelation loading and post-gelation loading. Pre-gelation loadinginvolves the incorporation of insulin solution into the polymer solutionprior to the gelation step. Proper dispersion is required to ensureuniform dispersion of insulin in the polymer solution. In contrast,post-gelation loading is achieved by equilibrating a solution containinginsulin with blank hydrogel particles. The small-sized insulin moleculescan diffuse into the semi-permeable hydrogel particles due to a concen-tration gradient.

5. Existing particle designs and their performance

Themethods described in the previous section are generally used tofabricate a core hydrogel structure. Thus, the simplest particle designsare beads or particles consisting of a single polymer. However, more

Fig. 3. Comparison of the size and size distribution of particles formed by different disper-sion techniques. Size distribution is indicated by the coefficient of variation (CV), where ahigh percentage of CV denotes a broad size distribution and vice versa. Note: The ranges ofCVs were calculated using the mean particle size and standard deviation reported in thereferenced works, where CV = (standard deviation/mean) × 100%.

Table 2Classification of particle designs and performance of biopolymeric oral insulin delivery systems.

Particle design (A) Material used (B) Size studied (C) Cumulativereleaseat acidicpHi, % (D)

Acid protection (E) Cumulativereleaseat alkalinepHj,% (F)

In vivos performance followingoral administration (G)

Ref.(H)

Chemicalstability

Retainedbiologicalactivity, %

Beads with a single polymer Alginate 40–150 nm 25 @pH 1.8

= n.d. N90b n.d. [75]

30–60 μm 0 to 8 n.d. 75 35 to 75b n.d. [64]53–112 μm 75 = n.d. 100 Blood glucose decreased to normal

level within 30 ming[52]

1.78 mm b8 n.d. n.d. 85d n.d. [73]Chitosan 315–1090 nm 63 n.d. n.d. 90b Sustained hypoglycemia for 6 h [59]

230–330 nm 20 to 90 n.d. n.d. 10 to 90d n.d. [53]265–387 nm 60 @

pH 4.0n.d. n.d. 85b,d Prolonged hypoglycemia for 15 h

and relativepharmacologicalbioavailabilityh(PA) was 15%

[74]

200–250 nm 55 @pH 1.0

n.d. n.d. n.d. Sustained hypoglycemia forapproximately 12 hand PA was approximately 3%

[76]

197 nm 28 @pH 2.5

n.d. n.d. 88b,d Sustained hypoglycemia for up to10 h

[77]

10 μm 63 n.d. n.d. 95b,e n.d. [57]Pectin 589–1000 nm 75 n.d. n.d. 50d n.d. [78]Gelatin 50–200 nm 55 @

pH 2.0= n.d. 100b,e n.d. [79]

Beads with blendedpolymers

Alginate–dextran sulfate 358–1067 nm 0 n.d. n.d. 93 n.d. [80]267–2760 nm 0 n.d. n.d. 100 n.d. [81]71 μm 0 n.d. n.d. 100 n.d. [82]

1.8 mm b8 n.d. n.d. 75 n.d. [73]Alginate–chitosan 1.5 mm 36 n.d. n.d. 70c n.d. [61]Alginate–whey protein 1.4–1.5 mm 70 n.d. n.d. 100d n.d. [83]

Coated beads Chitosan-coated alginate bead 748 nm 45 n.d. n.d. 70 PA was dose-dependent (rangingfrom 3.4 to6.8%) and hypoglycemic effectlasted for over 18 h

[68]

7.5 μm 5 n.d. n.d. 30 Rapid blood glucose reductionbegan at 12 h andwas sustained for more than 60 h

[71]

1.85 mm b8 n.d. n.d. 65d n.d. [73]1.06 mm 27 n.d. n.d. 20a,d n.d. [62]0.95 mm n.r. n.d. n.d. 70 n.d. [84]

Chitosan-coated pectinate bead 180–220 μm n.d. n.d. n.d. 5 to 60d n.d. [85]Chitosan-coated alginate–dextranbead

741–2652 nm 0 n.d. n.d. 65 n.d. [80]2.1 mm b5 n.d. n.d. 65d n.d. [73]

Albumin–chitosan-coatedalginate–dextran bead

110–392 nm 15 n.d. n.d. 75 Basal glucose level was reduced to48% at 12 h

[80]

Coated beadswith emulsifiedinsulin

Chitosan-coated nanoemulsion-loadedalginate beads

~488 nm 20 @pH 2.5

n.d. n.d. n.d. Blood glucose reduction wassustained for 10 h

[86]

88 @ pH 1Chitosan-coated lipoinsulin-loadedalginatebeads

0.8–1.2 mm n.d. n.d. n.d. 55b,d Blood glucose reduction wassustained up to 24 hand PA was dose-dependent andapproximately 7 to 8%

[72]

Colloidosomes Alginate gel-core with shells ofiron oxide nanoparticles

200–250 μm 90 n.d. n.d. 25b,d n.d. [87]

Alginate gel-core with shells ofsilica nanoparticles

~70 μm 70 n.d. n.d. 25b,d n.d. [88]

(continued on next page)

15H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

Table 2 (continued)

Particle design (A) Material used (B) Size studied (C) Cumulativereleaseat acidicpHi, % (D)

Acid protection (E) Cumulativereleaseat alkalinepHj,% (F)

In vivos performance followingoral administration (G)

Ref.(H)

Chemicalstability

Retainedbiologicalactivity, %

Hydrogel-coated particles Alginate-coated zinc calciumphosphate nano-particles

n.r. 9 n.d. n.d. 67d Sustained blood glucose reductionfor 12 h

[89]

Polyelectrolyte complexes Chitosan–alginate ~488 nm 100 n.d. n.d. n.d. n.d. [67]Chitosan–dextran sulfate ~500 nm 0 n.d. n.d. 80d n.d. [90]

~423 nm 50 n.d. n.d. 80d,f n.d. [67]

Polymer-insulin complexes Chitosan–insulin ~171 nm n.d. n.d. n.d. 30d n.d. [91]

Notes: n.d.: Not determined, n.r.: Not reported, =: Maintained.a pH of the incubation medium was 7.5.b pH of the incubation medium was 7.4.c pH of the incubation medium was 6.5.d Particles were not exposed to acidic medium prior to the alkaline testing.e Testing was performed for 3 h.f Testing was performed for 2 h.g Insulin released from microspheres in phosphate buffer saline was injected subcutaneously into diabetic rats.h Relative to subcutaneous injection of insulin.i Testing was performed for 2 h at pH 1.2.j Testing was performed for 4 h at pH 6.8 after i.

16 H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

sophisticated particle designs (see Table 2) can be developed with theaim of improving the stability and insulin delivery performance. Inthis section, the fabrication techniques andperformances for eachparticledesign are briefly described.

In addition, a summary of the insulin stabilization and deliveryperformances from the reported in vitro results for the various particledesigns developed using different biopolymers is given in Table 2. Theperformances of various particle structures and sizes were comparedon the basis of cumulative release of insulin over 2 h of incubation inacidicmedium (see ColumnD, Table 2); stability of insulin upon incuba-tion in acidic medium, as assessed by chemical and biological assays(see Column E, Table 2); and cumulative release of insulin in alkalinemediumafter acid-incubation (see Column F, Table 2). The in vivo efficacyof the oral-administered insulin-loaded biopolymeric particles was alsocompared.

Nevertheless, direct comparisons between the previous studies wasdifficult due to the varying materials and experimental conditionsadopted by different researchers, such as different material sources orcomposition, different hours of incubation, and different pH of theincubation medium. Thus, the comparison between results was donewith caution and was based on the best effort of the authors to identifythe general trends reported in the previous studies.

5.1. Beads with a single polymer

The most basic form of particle design in biopolymeric oral insulindelivery system consists of beads or particles made of a single polymer.The fabrication methods were briefly described in Section 4, and thesizes formed can range from 40 nm to 1.8 mm.

Fig. 5. Comparison of insulin release from particles of different sizes in alkaline medium.Notes: : Alginate beads; : Chitosan beads; : Alginate-dextran beads; : Chitosan-coat-ed alginate beads; : Chitosan-coated alginate-dextran beads.

Fig. 4.Comparison of insulin release from particles of different sizes in acidic medium.Notes: : Alginate beads; : Chitosan beads; : Alginate-dextran beads; : Chitosan-coated alginate beads; : Chitosan-coated alginate-dextran beads.

17H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

Alginate and chitosan are the two biopolymers mostly used in thisparticle design. In terms of performance, the release of insulin in acidicmediumwas generally lower from alginate beads compared to chitosanbeads (see Column D, Table 2). Both alginate and chitosan hydrogelsexhibit pH-sensitive characteristics, but in an opposite manner in acidicmedium. In an acidic environment, alginate hydrogels shrink due toincreased hydrogen bonding resulting from the protonation of thecarboxyl groups in the polymeric chains, which keeps insulin trappedwithin the hydrogel matrix [64,73]. Chitosan, on the other hand, swellsat acidic pH. As a result, a significant burst release of encapsulatedinsulin was observed when chitosan beads were incubated in acidicmedium [59,74]. In media of neutral or slightly alkaline pH, both alginateand chitosan beads swell, which promote the release of insulin, asobserved in Table 2 (column F). Other biopolymers such as pectin andgelatin have also been used to encapsulate insulin in the formof sphericalbeads, although the number of publications is scarce,with only one reportfor each of these biopolymers.

The protection of encapsulated insulin fromacid degradationhas beenevaluated in a few works but was generally limited to alginate beads.Degradation due to acidic pH [33,46,47,92] can disrupt the physicalstructure (i.e., physical stability), alter the chemical properties (i.e.,chemical stability), and lower the biological activity [4,50] of insulin.The chemical stability of retained insulin as determined using aUV-spectrophotometer [75] and high performance liquid chromatogra-phy (HPLC) [52] was reportedly maintained upon incubation in acidicmedium. Nevertheless, the biological activity of the retained insulin, asdetermined using a chemiluminescent immunoassay technique, wasreported to suffer losses of up to 25% upon incubation in acidic medium[64].

Encapsulation of insulin within biopolymeric matrices providesprotection against proteolytic degradation, as reported by Makhlofand co-workers [76]. Free insulin was completely degraded within5 min during acidic incubation in the presence of pepsin, whereasencapsulated insulin in chitosan-TPP nanospheres was protected fromenzymatic attack for at least 30min. The delayed proteolytic degradationof encapsulated insulin is most likely due to the presence of a polymericwall, which creates diffusion resistance against the inward penetrationof enzymes. When chitosan-TPP nanospheres were orally administeredto diabetic rats, the highest blood glucose reduction was achieved after4 h [76]. The hypoglycemic effect then reduced gradually and lasted forat least another 20 h.

5.2. Beads with blended polymers

Different biopolymers can be combined to form a blend of wallmaterials with enhanced physicochemical properties by exploiting thecharged properties of the polymers. The most frequently coupledpolymers in the development of biopolymeric oral insulin delivery

systems have been alginate and dextran sulfate. Other than dextransulfate, alginate polymer has also been coupled with polymers such aschitosan and whey protein. The incorporation of different materials intoa similar particle design can result in significantly different performance.

For instance, the performances of alginate–dextran sulfate beads inacidic medium were better than those of alginate–chitosan andalginate–wheyprotein,with almost insignificant insulin release occurring2 h of incubation [73,80–82]. This superior performance could be due tothe presence of the negatively-charged sulfate groups in dextran sulfatethat are able to prevent the release of the positively charged insulinmolecules at acidic pH. In contrast, the incorporation of whey protein,which swells at acidic pH, into alginate beads resulted in a burst releaseof approximately 70% of the encapsulated insulin during acidic incuba-tion. The performances of the blended polymer beads in alkalinemediumwere almost the same as the beads with single polymer discussed inSection 5.1. The cumulative release of insulin in alkaline medium after4 h of incubation ranged from 70 to 100% (see Column F, Table 2). Sta-bility tests assessing resistance to acidic and proteolytic degradationhave not been carried out for this form of particle design.

5.3. Coated beads

Acoating layer canbe applied to the beads fromSections 5.1 and5.2 toform coated beads. The coating process is normally carried out throughpolyelectrolyte complexation of oppositely charged biopolymers. Forinstance, chitosan is frequently used as a coating material for alginate oralginate-based beads because the cationic chitosan polymer is able tobind to the anionic alginate polymer almost instantaneously via electro-static interactions between the charged polymers. It should be notedthat the solution pH plays an important role in determining the chargedensities of both the alginate and chitosan polymers [93].

A second layer of coating composed of bovine serum albumin (BSA)has also been applied to chitosan-coated beads to act as a sacrificiallayer for proteolytic enzymes in the GIT [66,94]. BSA is a protein withan isoelectric point (pI) of 4.8 [95]. At pH values above its pI, BSAbecomes negatively charged and is, therefore, able to form a coatinglayer on positively charged chitosan surfaces due to electrostaticinteraction.

At acidic pH, the cumulative release of insulin from chitosan-coatedalginate beads ranged from 5 to 45%. While chitosan-coated alginate–dextran beads released less than 5% of their encapsulated insulin,albumin–chitosan-coated alginate dextran beads released approximately12 to 25% of their encapsulated insulin in acidic incubation medium. Theaddition of the albumin layer was claimed to weaken the interactionbetween insulin and the alginate beads [66]. The releases of insulin atalkaline pH were generally in the range of 65 to 70%.

Overall, the ranges of insulin release from the coated beads in bothacidic and alkaline media were observed to be lower than that fromthe non-coated beads. The presence of a coating layer surrounding the

18 H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

beads structure may have created a diffusion barrier that decreased thediffusivity of insulin across the shrunken or swollen beads. Moreover, thepresence of a coating layer successfully protected the encapsulatedinsulin from proteolytic degradation. The chemical stability of insulinreleased from chitosan coated-alginate microspheres incubated inacidic mediumwith pepsin was intact, whereas free insulin was totallydegraded [71].

Studies using in vivomodels revealed that the hypoglycemic effect ofcoated beads can generally be sustained longer compared to non-coatedbeads. For instance, when insulin-loaded and chitosan-coated alginatebeadswere administered to diabetic rats, the hypoglycemic effect lastedfor over 18 h [68]. The pharmacological bioavailability (PA) of the insulinrelative to a subcutaneous injection was dose-dependent and rangedfrom 3.4% to 7% [68]. In another instance, the hypoglycemic effect ondiabetic rats for oral-administered chitosan-coated alginate beads wassustained for more than 60 h, and their blood glucose levels slowlyreturned to the original level after 8 days [71].

5.4. Coated beads with emulsified insulin

The term ‘emulsified insulin’ represents the encapsulation of insulinwithin a hydrophobic core. In general, these particles designwere fabri-cated using a double emulsification technique, where emulsion solutioncontaining insulin encapsulated within a lipid or phospholipid wasdispersed into a biopolymer solution (i.e., alginate) that upon gelation,formed matrix-type structure and encapsulating the emulsified insulin.A coating layer was then applied using oppositely charged biopolymers(i.e., chitosan).

Although this particle designwas proposed to increase the encapsula-tion efficiency [86] and to improve the bioavailability of insulin duringtransit in the mucosal membrane of the intestinal tract [72], the incom-plete performance testing carried outmade it difficult to directly comparethe results between the published works. For instance, the release ofinsulin from chitosan-coated nanoemulsion-loaded alginate nanobeadswas almost 88% within 2 h of incubation in acidic medium with pH 1,but no similar testing has been carried out for chitosan-coatedlipoinsulin-loaded alginate beads. The high release of insulin fromchitosan-coated nanoemulsion-loaded alginate nanobeads was attrib-uted to the high acidity of the incubation medium destroying theintegrity of beads [72]. While alkaline testing was not performed forthe chitosan-coated nanoemulsion-loaded alginate nanobeads, approx-imately 55% of the encapsulated insulin was reported to be releasedfrom chitosan-coated lipoinsulin-loaded alginate beads after 4 h ofincubation in medium at pH 7.4 (see Column F, Table 2).

In vivo studies showed that the blood glucose reduction wasmaximized 6 h after the oral administration of chitosan-coatednanoemulsion-loaded alginate beads in diabetic rats, and the bloodglucose level reached normal values 10 h after administration [86]. Forlipoinsulin-loaded alginate beads, the blood glucose profiles of diabeticrats were more encouraging, as blood glucose reduction was sustainedfor up to 24 h [72]. This prolonged sustained release could be attributedto themucoadhesive properties of chitosan and alginate on the intestinalmucosa that prolonging the retention time in the GIT.

5.5. Colloidosomes

Colloidosomes are microspheres with shells composed of closelypacked colloidal particles formed by a Pickering emulsion [96]. Thepermeability and pore size of the colloidal shell can be varied accordingto the type, size, and degree of fusion of the colloidal particles [97,98]that are used to stabilize the dispersed droplets in the emulsion [99].Examples of colloidal particles that have been used in the developmentof biopolymeric colloidosomes include magnetic iron oxide and silicananoparticles [87,88].

Initially dispersed in the oil phase, the colloidal particles will adhereto the surface of alginate droplets within the emulsion system through

self-assembly at the liquid–liquid interface. The gelation of the alginatedroplets has been achieved by the hydrolysis of D-glucono-δ-lactone(GDL), pre-mixed in the alginate solution, into an acidic solution thatcauses the release of calcium ions from a Ca-chelating agent, such ascalcium-ethylenediaminetetraacetic acid (Ca-EDTA), also pre-mixed inthe alginate polymer solution.

Observation of the release profiles from the previous works [87,88]revealed that approximately 70 to 90% of the encapsulated insulin wasreleased in acidic medium within 2 h of incubation. This high releaseof insulin was attributed to undissolved insulin microcrystals accumulat-ed at the surfaces of the colloidosomes [88]. Moreover, the porous shellsof the colloidosomes, which are due to the interstitial pores formedbetween the colloidal particles, could have also led to the high releasesof insulin. At alkaline pH, the presence of colloidal shells hampered theswelling of the alginate gel-core, resulting in only approximately 25%release after 4 h of incubation in alkaline medium.

5.6. Hydrogel-coated particles

Other than forming the core structure for insulin encapsulation,alginate can also be used as a coating material. Alginate-coated zinccalciumphosphate nanoparticles have been developed to act as a carriersystem for oral insulin. The diffusion of free calcium ions from theinsulin-loaded zincmodified calcium phosphate nanoparticles immersedin alginate solution triggered the gelation of the alginate polymer, thusforming a coating layer on the nanoparticles [89].

The in vitro release study of the alginate-coated particles showed lessthan 9% release of the encapsulated insulin into the acidic medium [89].Shrinking of the alginate layer in acidic medium provided a diffusionbarrier to the large molecular weight-hexameric insulin (MW = 32–34 kDa) used in the study. In contrast, the release in the alkalinemedium was approximately 67%. Nevertheless, the stability of theinsulin retained from acidic incubation and incubation in the presenceof proteolytic enzymes was not evaluated. Oral administration of thehydrogel-coated particles in diabetic rats enabled sustained reductionof blood glucose levels for almost 12 h [89].

5.7. Polyelectrolyte complexes

Polyelectrolyte complexes were formed when two oppositelycharged polymer solutions solidified to form particles upon pH adjust-ment. Chitosan-alginate and chitosan-dextran sulfate complexes werefabricated through drop-wise addition of the polycationic solution(i.e., chitosan) into the polyanionic solution (i.e., alginate or dextransulfate) that was stirring at a pre-determined pH. The resulting com-plexes entrapped insulin within the polymer networks.

Depending on the polyanion used, the amount of released insulin inacidic medium varied from insignificant (i.e., 0%) to full (i.e., 100%)release. For instance, chitosan–alginate complexes gave full release ofinsulin into the surrounding acidic medium due to the weakenedintegrity of the complexes at acidic pH [67]. Chitosan–dextran sulfatecomplexes, however, were successful in fully retaining the insulin atacidic pH due to the presence of the negatively charged sulfate groups[90]. At an alkaline pH of 6.8, the weakened ionic interactions betweenboth the negatively charged insulin and the sulfate groups resulted inapproximately 80% of encapsulated insulin being released into theincubation medium [67].

5.8. Polymer-insulin complexes

Depending on its source, the pI of insulin varies from5.5 to 6.4. At a pHabove the pI value, insulin becomes negatively charged and vice versa.This characteristic of insulin was utilized to form polymer-insulin com-plexes through electrostatic interactions with a cationic polymer, namelychitosan, which is protonated at pH values below its pKa value of 6.5. Inbrief, the pH of a chitosan solution was adjusted to 5.5 to keep it

19H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

protonated, whereas an insulin solutionwas adjusted to pH 8 to render itnegatively charged [91].When both solutions weremixed under stirring,complexes were formed almost instantaneously. The performance ofthese particles in acidicmediumhas not been tested; in alkalinemedium,only 30% of the insulin was released.

5.9. Effect of particle size

Based on the data summarized in Table 2, a further evaluation on thegeneral trend for the effect of particle size on the release of insulin inincubation medium was carried out. To facilitate unbiased judgment,only particle designs for which different size ranges have been reportedwere considered. These included beads made of alginate and chitosan,alginate–dextran, chitosan-coated alginate beads, and chitosan-coatedalginate–dextran beads (see Figs. 4 and 5).

In acidicmedium (see Fig. 4), the particle designswith larger sizes ofmicro- to millimeter in diameter tend to give lower cumulative releaseof the encapsulated insulin. This is especially true in the case of alginatebeads, chitosan beads, and chitosan-coated alginate beads. High releasesof approximately 45 to 90% were often reported for smaller particles ofnano- to micrometers in diameter (see Fig. 4). In smaller beads, thediffusion length for the drug to reach the outer medium is shorter[100] compared to larger beads. Moreover, the high release insmaller beads could be due to the high concentration of insulinaccumulated at the surfaces of the beads [87,88]. Additionally, the higherspecific surface area and the shorter diffusion length in smaller beadsoften led to a burst release of encapsulated drug during the incubationtesting.

For the alginate–dextran beads and the chitosan-coated alginate–dextran beads, an effect of particle size on the release of insulin inacidic medium is not seen. Both millimeter and nanometer-sizedbeads showed insignificant release of insulin in acidic medium.This suppressed-release of insulin from both large and small beadscould be attributed to the interaction between the sulfate groups ofdextran sulfate with the insulin according to the electrostatic interactionmechanism described earlier.

In alkaline medium, all of the aforementioned particle designsshowed no obvious effect of particle size. Generally, all the particledesigns mentioned were able to release approximately 70 to 90% ofthe encapsulated insulin in alkaline medium (see Fig. 5).

6. Concluding remarks and future perspectives

Polymeric encapsulation is one of the most employed approachesused by researchers in attempts to improve the bioavailability of insulinupon oral administration. Biopolymers, which are hydrophilic in general,are favorable materials for encapsulation of insulin, which is also hydro-philic. Importantly, the biopolymers and particle designs used are criticalin determining the performance of the encapsulation system in terms ofprotecting the insulin from acid degradation, reducing the leakage ofinsulin in acidic medium, and releasing the encapsulated insulin inalkalinemedium. In general, alginate is themost widely used biopolymerto form the particle core. Alginate shows better performance in reducingthe leakage of insulin in acidic medium compared to chitosan, due to itsstability in acidic medium. The ability of sulfate groups incorporatedwithin polymer matrices to reduce the leakage of insulin in acidicmedium but release the insulin in alkaline medium is also noteworthy.Moreover, coating reduced the leakage of insulin in acidic medium anddelayed the release of insulin in alkaline medium due to the extradiffusional barrier surrounding the beads. The latter effect can be utilizedas a potential strategy to sustain the release of insulin at the targeted sitein the intestinal tract. Other particle designs that have been recentlydeveloped, such as polyelectrolyte complexes and polymer-insulincomplexes, are promising but require further studies to improve theirperformance.

Biopolymers are derived from natural sources. Thus, their composi-tions are subject to seasonal variation and also depend on the speciesof plants or animals from which they are derived. The effect of suchvariations on the stabilization and controlled-delivery of insulin needsto be evaluated in the future. However, these biopolymers can be chem-ically modified to improve their stabilization and delivery performance.For example, chitosan has been chemically modified to improve itsstability in acidic medium [101]. Hydrophobic modification of chitosanby acylation with fatty acids reduced the leakage of insulin in acidicmedium by selecting the fatty acids with longer alkyl chains [57,101].The incorporation of hydrophilic and hydrophobic moieties (i.e., succinyland lauryl group, respectively) to chitosan further suppressed the releaseof insulin in acidicmedium to only approximately 8.5% [59]. However, themodification process often involves the use of hydrocarbon-basedchemicals such as acyl chlorides [57,101] and organic solvents such asmethanol [57,101], which may be toxic to humans and are potentiallyharmful to the environment. Future work should be directed towardusing chemicals and processes that are human and environmentallyfriendly.

Ideally, upon oral administration of insulin-loaded particles, the bloodglucose level will begin to reduce gradually after 1 to 2 h as the ingestedparticles reach the intestinal area, which is the targeted site of absorption.The hypoglycemic effect induced by insulin released thereafter should besustained for a long period of time without significant peaks in activity.Sustained release of biologically active insulin is important to reducethe frequency of medication required by the diabetes patient. Generally,the in vivo performance of insulin-loaded biopolymeric oral insulindelivery systems have shown promising results, as most of the particledesigns tested have been able to sustain a hypoglycemic effect for atleast 6 to 12 h. The hypoglycemic effect may be further prolonged forup to 60 h by coating the beads [71]. Thus, it is possible to tune thein vivo performance of oral-administered insulin by manipulating theparticle designs.

However, the pharmacological bioavailabilities of oral-administeredinsulin-loaded biopolymeric particles relative to subcutaneous injec-tions of insulin were generally found to be low, ranging from 3% to15% [74,76]. This low bioavailability could be partly due to the degrada-tion of insulin by proteolytic enzymes and acidic pH. This speculation issupported by the reduction in bioactivity level of insulin upon exposureto acidic medium [64] and the loss of chemical integrity when insulin isexposed to enzymes [71,76,80] in vitro. Nevertheless, the number ofstudies that have investigated the in vitro bioactivity of insulin upon in-cubation in acidic or alkalinemedia is still limited. Previous studies havebeen more focused on the effect of encapsulation on the sustainedrelease profile of insulin in simulated intestinal fluid. Moreover, thein vitro testing of insulin stability requires the use of an immunoas-say technique, which is costly and requires a specific analyticalinstrument. Future work should seek to understand how the encap-sulation and different particle designs are able to improve insulinbioactivity.

In conclusion, encapsulation of insulin within matrices of bio-polymers is a promising approach for the oral delivery of insulin,as shown by the encouraging in vitro and in vivo results reportedto date. Biopolymers are inexpensive, safe and biocompatible, mak-ing them attractive from both economic and safety viewpoints. Inaddition, the encapsulation methods can be performed under mildconditions (i.e., room temperature and near neutral pH), makingthem suitable for the encapsulation of insulin and other proteindrugs in general, compared to other protein encapsulation systems.However, more research is still required to improve and confirm thebioavailability, stability, and release profiles of insulin using in vitrosystems and in vivo animal models before these delivery systemscan move to the next phase of clinical trials and then commerciali-zation. This review provides a basis for future work on the optimiza-tion of particle design and material formulation to achieve therequired improvements.

20 H.-P. Lim et al. / Journal of Controlled Release 186 (2014) 11–21

Acknowledgments

The authors wish to thank the Ministry of Science, Technology andInnovation of Malaysia for supporting this research under the E-Sciencegrant (06-02-10-SF0157) and the School of Engineering, MonashUniversity Malaysia for providing the Ph.D. studentship for Lim HuiPeng.

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