Controlled Delivery of Basal Insulin from Phase-SensitivePolymeric Systems After Subcutaneous Administration:In Vitro Release, Stability, Biocompatibility, In Vivo Absorption,and Bioactivity of Insulin
KHALED AL-TAHAMI, MAYURA OAK, JAGDISH SINGH
Department of Pharmaceutical Sciences, College of Pharmacy, Nursing, and Allied Sciences, North Dakota State University,Fargo, North Dakota 58105
Received 18 August 2010; revised 7 November 2010; accepted 12 November 2010
Published online 9 December 2010 in Wiley Online Library (wileyonlinelibrary.com). DOI 10.1002/jps.22433
In healthy individuals, there are two types of insulinsecretions: basal and stimulated. Basal insulin is se-creted between meals and throughout the night ata constant rate of 0.5–1 U/h to sustain serum con-centrations of 5–15 :U/mL. On the contrary, stim-ulated insulin secretion occurs in response to mealsand results in insulin concentrations of 60–80 :U/mL 30 min after the meal.1 Although the basal in-sulin level is low, it modulates hepatic glucose produc-
Correspondence to: Jagdish Singh (Telephone: 701-231-7943;Fax: 701-231-8333; E-mail: [email protected])
tion rate and glucose output during prolonged inter-vals between meals. The importance of basal insulinin forestalling or at least postponing the long-termcomplications of diabetes has been well established.2
The Diabetes Control and Complications Trial endeddecades of controversy of whether maintaining nor-mal blood glucose could prevent or postpone diabetescomplications.3 It has been shown that intensive in-sulin therapy of type I diabetic patients to eliminatethe fluctuation in basal insulin levels reduced the riskof retinopathy by 22–54% and risk of nephropathy by60%.
Because of extensive degradation of insulin whenadministered orally, it is currently delivered via par-enteral route. Insulin therapy involves one or moredaily doses of intermediate- or long-acting insulininjection to satisfy basal insulin requirement. Thisregimen affects patients’ life style and leads to poor
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compliance, pain, and mental stress. Therefore, thereis a need to develop controlled-release delivery sys-tem to deliver basal insulin for longer duration aftera single subcutaneous injection.
(Genentech, South San Fransisco, CA, USA), for thecontrolled delivery of growth hormone. This regula-tory approval for biomedical use in humans triggeredfurther exploration of these delivery systems for otherproteins, including insulin. Injectable microparticlesfrom poly(D,L-lactide) (PLA) and poly(D,L-lactide-co-glycolide) (PLGA) have been investigated for the de-livery of insulin.5–9 A high initial burst release wasobserved followed by a slow release over 1–2 weeks pe-riod. The release profile was affected by the additionof hydrophilic additives such as water and glycerol.Microspheres composed of alginate have also been ex-plored for insulin delivery but the release did not lastmore than a day.10 Microspheres prepared from chi-tosan showed a biphasic release profile characterizedby a high initial release.11 In addition, microspheressuffer several inherent disadvantages including therelatively complicated manufacturing procedure, in-sulin stability, and low drug loading capacity.
The stability of a protein is a major concern be-cause physically and chemically modified proteinscarry the risk of losing biological activity. Insulin ag-gregation is also a major physical instability. Deami-dation, cleavage of peptide bonds, and hydrolysis arealso noted instabilities of insulin.12 Therefore, confor-mational, secondary structure and chemical stabili-ties are important parameters to be investigated forinsulin.
One of the key factors in developing a controlled-release system for commercial use is its ability tocontrol the release rate in order to attain and main-tain the desired therapeutic level of the drug. Ideally,a zero-order drug release from the delivery systemis desired because it assures constant drug releaseregardless of concentration. However, such a releasepattern is not always feasible. Many delivery systemsshow high initial burst release, and as the amountof drug within the delivery system diminishes, its re-lease rate decreases. The release of drugs from poly-meric devices occurs through diffusion of the incorpo-
rated drug, polymer erosion, or a combination of bothmechanisms.13
Phase-sensitive polymeric formulations have beeninvestigated for the controlled release of variousproteins such as lysozyme and human growthhormone.14,15 One of the problems is the burst re-lease during the first few hours following injectioninto the body. This could be due to the lag time be-tween the injection of the delivery system and theformation of the gel depot. Burst release of proteinswas modulated in benzyl benzoate/benzyl alcohol sol-vent systems by controlling the composition of thesolvent system and polymer concentration.14,16 In-creasing the proportion of the hydrophilic solvent ledto an increase in the burst release. Another studyshowed that the release of proteins can be modulatedby using different polymer concentrations, alteringthe solvent composition, and using proteins of differ-ent molecular weights (MWs).17
At low concentrations, insulin is present in themonomeric form. At higher concentrations, insulinmonomers associate to form dimers and, if zinc ionsare present, they form hexamers.18 Insulin dimericstructure is formed by nonpolar forces and hydrogenbonding. The association of the dimers to form hexam-ers around two zinc ions is associated with the burialof remaining nonpolar groups.19 The stability of in-sulin is highly affected by the association state, withthe hexamer form being the most stable. Zinc ionsexert their stabilizing effect by neutralizing negativecharges in the center of the insulin hexamer. In addi-tion, insulin hexamer is larger in size and thereforewould diffuse slowly from the delivery system in orderto prolong the release and reduce the burst release ofinsulin.
The in vitro biocompatibility of polymeric deliverysystems is essential for their utility in clinical appli-cations. A number of PLA-containing systems havebeen approved by the FDA for drug delivery.20 How-ever, the response of tissue to polymeric implants de-pends on many factors including amount of polymer,nature of degradation products, rate of degradation,and degradation product accumulation in the tissue.Therefore, evaluation for polymeric delivery systemsbiocompatibility in vitro is indispensable.
In a recent study, injectable pentablock copoly-mer hydrogels composed of poly($-amino ester)-poly(g-caprolactone)-poly(ethylene glycol)-poly(g-caprolactone)-poly($-amino ester) were evaluated forbiocompatibility and sustained release.21 These pH-and temperature-sensitive polymeric system werefound to be biodegradable and provided a sustainedrelease of insulin for 15 days.
In this study, phase-sensitive system (PLA in tri-acetin) was investigated for controlled release ofinsulin in vitro and In Vivo in diabetic rats. Theconformational and chemical stability of the in vitro
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released insulin was evaluated using differentialscanning calorimetry (DSC), circular dichroism (CD),and matrix-assisted laser desorption/ionization time-of-flight (MALDI-TOF) mass spectrometry. Finally,biocompatibility of phase-sensitive polymer deliv-ery system was evaluated in vitro using 3-(4,5-dimethylthiazol-2-yl)-2, 5-diphenyltetrazolium bro-mide (MTT) assay. In Vivo biocompatibility of the de-livery system was evaluated by analyzing the tissuesamples using light microscopy.
EXPERIMENTAL SECTION
Materials
D,L-Polylactide [inherent viscosity 0.2 dL/g, MW≈ 10,000) was obtained from Polysciences Inc.(Warrington, Pennsylvania). Human recombinant in-sulin was purchased from Celliance Corporation(Norcross, Georgia). Triacetin and dimethylthiazolyldiphenyltetrazolium bromide were procured fromSigma Chemical Company (St. Louis, Missouri). Mi-cro BCA protein assay kit was purchased from PierceBiotechnology Inc. (Rockford, Illinois). Human insulinELISA kit was purchased from Linco Research Inc.(St. Charles, Missouri). Human embryonic kidney(HEK293) cell line was obtained from American TypeCulture Collection (ATCC, Rockville, Maryland). Allother chemicals used were of analytical grade.
Preparation of Phase-sensitive Delivery Systems and InVitro Release
as a blank control for absorbance. The amount of in-sulin released in the samples was obtained from thestandard curve and corrected for sample removal.23
Secondary structure of the released insulin, as well asof insulin solution, was evaluated using CD. Spectraof released insulin in the samples were recorded ona Jasco J-815 CD Spectrophotometer (Jasco, Tokyo,Japan). The filtered samples were scanned in a cellwith 1 mm path length from 200 to 300 nm, usinga bandwidth of 1 nm, 0.5 s response time, 0.2 nmdata pitch, and a scanning speed of 100 nm/min. Eachspectrum is the average of three scans. Spectra ofthe buffer were recorded and subtracted from sam-ple spectra. Insulin concentrations obtained by BCAassay were used to calculate the molar ellipticity (θ).
Mass Spectrometry
MALDI-TOF mass spectrometry was used to in-vestigate the chemical stability of released insulinfrom polymeric delivery systems. Matrix solutionwas prepared by dissolving 10 mg of "-cyano-4-hydroxycinnamic acid in 1 mL mixture containing 1:1ratio of acetonitrile and 0.1% (v/v) Trifluoroacetic acid(TFA) water solution. Ten microliters of insulin sam-ple was added to 100 :L of matrix solution and mixedby vortexing. An aliquot (2 :L) of the final solutionwas applied to the sample target plate and allowed todry prior to analysis. MALDI-TOF experiments werecarried out on a Bruker MALDI TOF II (Bruker Dal-tonics Inc., Billercia, Massachusetts) equipped witha 200 Hz solid-state smartbeam laser. Samples wererun in the positive reflectron mode and data were
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Male Sprague–Dawley rats, 8 weeks old and weigh-ing 220–224 g, were purchased from Zivic Laborato-ries, Inc. (Pittsburg, Pennsylvania). All animal exper-iments were conducted in compliance with humaneanimal care standards outlined in the “National In-stitutes of Health (NIH) Guide for the Care and Useof Experimental Animals” and according to the pro-tocol approved by the Institutional Animal Care andUse Committee at North Dakota State University. Di-abetes was induced by a single intraperitoneal injec-tion of streptozotocin (STZ; 60 mg/kg) dissolved in cit-rate buffer (pH 4.5). For the first 24 h after injection,rats were provided with 5% sucrose solution insteadof water to counteract hypoglycemia, which may occurdue to $-cell necrosis and endogenous insulin release.The rats were considered diabetic if fasting blood glu-cose level was higher than 200 mg/dL 1 week afterSTZ injection.
The in vitro biocompatibility of phase-sensitive de-livery systems was evaluated for the effect on mito-chondrial succinate dehydrogenase activity by MTTassay.25 PLA–triacetin phase-sensitive systems wereextracted in PBS (pH 7.4) by keeping samples of500 :L gel depot for 10 days at 37◦C and 70◦C, re-spectively. Medium prepared by the same way withoutpolymer addition served as negative control, whereas2% dimethyl sulfoxide (DMSO) in growth mediumserved as positive control.25 After 10 days, the pH ofextracts was measured and adjusted to 7.4 by adding1 M NaOH. The extracts were then filtered and di-
software.The In Vivo biocompatibility of the polymeric de-
livery systems was evaluated by studying rat’s skintissue for inflammatory changes after administrationof the polymeric delivery systems. The study involvesinjecting 100 :L of formulation subcutaneously intothe upper portion of neck of the rat where a visiblegel lump was formed. At different time points (days 1,7, 30, and 90), the rats were euthanized and skin tis-sue from injection sites were surgically removed. Skinsamples were then fixed in 10% neutrally bufferedformalin solution. After collection of the skin sam-ples, they were washed with water to remove ex-cess fixative, dehydrated by transferring to increas-ing strengths of alcohol, and embedded in paraffin.Transverse sections of 5 :m thick were cut by rota-tory microtome, mounted on a glass slide, and stainedwith hematoxylin and eosin. The slides were observedunder light microscope for the presence of any signsof acute and chronic inflammation, fibrous capsuleformation, fibrosis, tissue morphology, and necrosis.
Data Analysis
The results are presented as mean ± standard devi-ation. Statistical comparisons were carried out usinganalysis of variance and Student’s t-test. The proba-bility (p value) of less than 0.05 was considered to besignificant.
The bioavailability and pharmacokinetic evalua-tion was based on the area under the curve (AUC)calculations. Area under serum insulin concentrationversus time was estimated by trapezoidal rule.26 Thefollowing equations were used to calculate AUC:
AUC0 − n =n∑
i=0
{Ci + Ci + 1
2• (ti + 1 − ti)
}(1)
AUCn − ∞ = ClastKel
(2)
where Ci is serum insulin concentration at intervali, n is number of sampling time intervals, Clast is theconcentration at last time point, and Kel is elimination
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rate constant calculated from the slope of the finalsegment of serum concentration versus time curve.
RESULTS AND DISCUSSION
In Vitro Release
In vitro release studies were carried out to evalu-ate the effect of varying divalent metal ions, poly-mer concentration, and insulin loading dose on therelease and release kinetics from PLA-based systems(Table 1). The first long-acting insulin was developedin 1936 by complexing insulin with protamine andzinc. Ever since, commercially available insulin hasalways contained zinc ions. The addition of zinc saltsresults in the association of insulin with hexamers,which has lower solubility, and therefore longer ac-tivity as absorption is impeded.27
The formation of the insulin hexamer upon the ad-dition of zinc ions was confirmed by a series of DSCstudies (data not shown). As the molar ratio of zincion to insulin is increased, the DSC monomer tran-sition peak (Tm ∼ 68◦C) was attenuated with an in-crease in the hexamer peak (Tm ∼ 82◦C). At a ratio of1:5 insulin hexamer to zinc ion, insulin showed onlythe hexamer peak denoting the presence of insulincompletely in the hexamer form. Similar results werereported in the study by Huus et al.,28 wherein atzinc ratio of at least 5, only the hexamer transitionpeak was observed. In an attempt to improve insulinrelease profile, hexameric insulin formed with zincacetate was incorporated into 30% PLA–triacetin de-livery system. Figure 1 shows the effect of differentzinc ion concentrations on the release profile of in-sulin. The initial burst release was reduced when zincion was added to the delivery systems. Introductionof high amounts of metal ion (1:50 insulin hexamer to
Figure 1. In vitro release of insulin from in situ gel depots(effect of metal ion content) (n = 4). �, No zinc added; •, 1:5insulin hexamer to zinc ion; �, 1:50 insulin hexamer to zincion; ∗, significant compared with 1:5 insulin hexamer to zincion; and #, significant compared with 1:50 insulin hexamerto zinc ion at p value < 0.05.
zinc ion ratio) caused a higher release rate comparedwith delivery systems containing lower zinc ion con-tent (1:5 ratio). All delivery systems showed a best-fitfor Higuchi model, which describes drug release fol-lows a diffusion process and is dependent on squareroot of time. The delivery systems prepared with a 1:5insulin hexamer to zinc ion ratio showed higher cor-relation with Higuchi model and zero-order releasekinetics than other systems that contain no or higheramounts of zinc (Table 1). Consequently, the ratio of1:5 was adopted for further insulin delivery studies.
Insulin monomer (MW ≈ 5808 Da) is highly solublein aqueous solutions and it can easily diffuse throughthe polymeric porous matrix due to its low MW. Asso-ciation of insulin with hexamers reduces its solubilityin water, which reduces insulin dissolution and waterpenetration into the depot and, consequently, insulinrelease. The higher insulin release from delivery sys-tems containing high amount of zinc acetate could beattributed to the high water solubility of zinc acetate,which could lead to higher water penetration into thedepot.
Figure 2 shows the effect of PLA concentration (dis-solved in triacetin) on insulin release. Delivery sys-tems containing 20% PLA (w/v) showed higher burstrelease (16.29%) and the insulin was released within 2weeks, whereas delivery systems containing 30% and40% PLA showed an extended release of insulin overperiods of 4 and 8 weeks, respectively. All deliverysystems showed best fit for Higuchi model followedby zero-order release kinetics (Table 1). Increasingthe polymer concentration has been shown to slowthe phase separation rate and decrease the waterinflux.29 Lower rates of water influx and liquid–liq-uid phase separation at higher polymer concentra-tions have also been reported.30 Increasing the poly-mer concentration results in a multitude of effects,
Figure 2. Effect of PLA concentration on the in vitro re-lease of insulin from PLA–triacetin phase-sensitive deliverysystems containing 1:5 insulin hexamer to zinc ion molarratio (n = 4) (effect of PLA concentration). �, 20% PLA; •,30% PLA; and �, 40% PLA.
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Table 1. Composition of Phase-Sensitive Insulin Delivery Systems and In Vitro Release Kinetics
Formulation D,L-PLA (% w/v) Triacetin (% v/v) Insulin (% w/v)Insulin Hexamer/Metal Ion Ratio Zero-Order r2 First-Order r2 Higuchi r2
Release kinetics equations: zero-order (Mt/M∞ = kt), first-order [ln(Mt/M∞) = kt], and Higuchi (Mt/M∞ = kt0.5), where Mt/M∞ is the fractional drugrelease, t is the release time, and k is a kinetic constant.
including lower system diffusivities and increased so-lution hydrophobicity, all of which may contribute tothe dynamics of water influx reduction. The slowerrelease rate implies a slower degradation rate of thegel depot.
Finally, the effect of insulin loading on the in vitrorelease from 30% PLA–triacetin systems was investi-gated. Delivery systems that contain the higher dose(25 mg) exhibited a greater initial burst release ofinsulin (Fig. 3a). This could be due to the fact thatincreasing insulin loading leads to the formation ofgel depots with higher amounts of insulin on thesurface. Once a gel is formed in the release media,insulin molecules on depot surface dissolve, leavingmore pores compared with delivery systems contain-ing less insulin. These porous structures lead to fur-ther drug release and degradation of the matrix. Sim-ilar results were observed when PLGA–acetonitrilesystems were used for the delivery of amylase31 andgrafted PLGA–glucose microspheres for the deliveryof octreotide.32 Both formulations exhibited high cor-relation for the Higuchi diffusion model (Table 1 andFig. 3b).
Stability Studies
Many destabilizing factors could affect the physicaland chemical stability of insulin within the deliverysystems. Hence, it was extremely crucial to investi-gate the stability of the released insulin from phase-sensitive delivery systems. The conformational stabil-ity of the released insulin from delivery system wasstudied by DSC, which showed that the insulin re-leased at day 14 had lower �H and Tm compared withcontrol, which is fresh insulin (Fig. 4). Fresh and re-leased insulin showed �H values of 13.4 ± 0.4 × 103
and 8.1 ± 0.3 × 103 cal/mol and Tm values of 82.77 ±0.64 and 81.05 ± 0.41◦C, respectively.
Figure 5 represents CD spectra for insulin released,fresh insulin, and insulin solution at day 14. Therewas a reduction in the magnitude of both minimaat 209 and 222 nm in released insulin as comparedwith fresh insulin. The secondary structure of in-sulin solution incubated for the 14 days was com-pletely destroyed as determined by CD spectroscopy.
The partial reduction in conformational stability andsecondary structure is common in released peptideand protein samples.33 Many factors such as temper-ature and agitation could have contributed to reduc-tion in the released insulin stability.34 The reductionin released insulin conformational stability could beattributed to insulin prolonged existence in agitatedmedia at 37◦C temperature.34 We do not expect thisto happen in the In Vivo studies as released insulin
Figure 3. In vitro release (a) and Higuchi model fitting (b)of insulin from 30% PLA–triacetin in situ gel depots withdifferent insulin loading dose (n = 4). Insulin concentrationof 1% (w/v; �) and 5% (w/v; �).
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Figure 4. DSC scans of released insulin at day 14 fromphase-sensitive delivery system [30% PLA–triacetin and5% (w/v) insulin]. Fresh insulin (solid line), released insulin(dashed line), and insulin solution (dotted line).
will be rapidly absorbed. Yet, conservation of most ofdenaturation peak in DSC thermogram and CD char-acteristic spectrum demonstrate that the conforma-tional stability and secondary structure were greatlypreserved during the delivery systems preparationand release.
Insulin is susceptible to several chemical modifi-cations, which would alter its primary structure andpossibly its biological activity. The chemical stabilityof released insulin from phase-sensitive delivery sys-tems was evaluated using MALDI-TOF mass spec-troscopy. Figure 6 shows the MALDI spectrum of2-week released insulin sample from phase-sensitivedelivery system. The insulin integrity was conservedin the released samples with (M+H)+ signal corre-sponding to a molecular mass of 5808.8 Da. Therewas no significant degradation products formed dueto peptide bonds breakage. Acylation products (e.g.,lactyl insulin derivatives) are usually characterized
Figure 5. CD spectra of released insulin at day 14 fromphase-sensitive delivery system [30% PLA–triacetin and5% (w/v) insulin]. Fresh insulin (solid line), released insulin(dotted line), and insulin solution (dashed line).
Figure 6. MALDI-TOF mass spectroscopy of released in-sulin (day 14). Insert shows the enlarged spectrum exhibit-ing a peak corresponding to 5792.3 Da, indicating the lossof a water molecule, possibly due to the formation of a cyclicimide product.
by the presence of signals in the region (MW + 72).We did not observe any acylation products signalsor signals corresponding to high MW transformationproducts such as the formation of covalent insulindimers. Nevertheless, released insulin from phase-sensitive delivery system showed a signal correspond-ing to a MW of 5792.7 Da. This indicates a loss ofa water molecule from the original insulin, whichsuggests the formation of cyclic imide product. Thisreaction usually follows deamidation reactions andmay lead to the formation of isoAsp-insulin.35 Insulinundergoes many chemical modification reactions de-pending on the surrounding environment. Yet, theeffect on biological activity varies depending on thenature of the forming products. Deamidation prod-ucts were shown to have almost the same biologicalactivity as native insulin, whereas covalent insulindimers and hydrolysis products (A8–A9) were shownto exhibit only 15% and 2% relative potency,respectively.12
In Vivo Absorption
The ultimate success of the delivery systems for thecontrolled release of insulin at basal level is their abil-ity to control the In Vivo release and lower bloodglucose level for longer duration. Figure 7 showsthe serum insulin and blood glucose levels followingsingle subcutaneous administration of insulin-loadedphase-sensitive delivery system (30% PLA dissolvedin triacetin) and insulin solution in PBS. Three groups(six animals in each group) received phase-sensitivedelivery systems with no insulin (control), insulinalone (60 U/kg body weight), and insulin (60 U/kgbody weight) and zinc acetate in the molar ratio of1:5 insulin hexamer to zinc ion ratio. A solution group
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Figure 7. Serum insulin concentration (a) and blood glucose level (b) following subcutaneousadministration of insulin solution, serum insulin concentration (c), and blood glucose level(d) following subcutaneous administration of phase-sensitive polymeric delivery systems inSprague–Dawley rats (n = 6). �, Blank phase-sensitive delivery systems (control); �, phase-sensitive delivery systems containing insulin; and •, phase-sensitive delivery systems contain-ing zinc–insulin.
was studied by injecting subcutaneously insulin dis-solved in PBS at a dose of 2 U/kg body weight. Seruminsulin level increased rapidly, reaching mean peakconcentration (Cmax) of 67.84 :U/mL at 2 h after ad-ministration and declined afterwards to reach below2 :U/mL (detection limit) after 12 h in the solutiongroup (Fig. 7a). Blood glucose levels decreased fol-lowing insulin absorption, showing an acute and rel-atively short hypoglycemic effect. Blood glucose lev-els were restored to pre-administration levels within6–8 h (Fig. 7b). Treatment groups showed contin-uous release of insulin from phase-sensitive deliv-ery systems over extended periods of time. The AUCvalues were 14,582.52 ± 2067.51 and 34,832.13 ±5879.19 :U·h/mL for the treatment groups receivinginsulin alone and zinc–insulin, respectively (Figs. 7cand 7d). The delivery systems containing insulin andzinc salt showed significantly higher (p < 0.05) seruminsulin level over a longer period compared with con-trol group. The released insulin exhibited fluctuationsfollowing administration. Yet, basal insulin levels arelow and a slightly higher initial release should notimpose a major concern. In addition, fluctuation ofserum basal insulin levels has been reported with ex-isting insulin therapies such as Neutral ProtamineHagedorn (NPH) insulin and insulin ultralente.36 Asignificant (p < 0.05) reduction in blood glucose levels
was observed over 4 weeks following the administra-tion of delivery systems containing insulin and zincsalt.
Addition of zinc, as discussed earlier, decreasesinsulin solubility, which impedes its release. Addi-tionally, insulin in the delivery system, which con-tained no zinc, may have undergone denaturationand consequently aggregation that would lower in-sulin levels in blood. Denaturation is usually ob-served under circumstances that favor the formationof insulin monomers.37 Therefore, fibril formation isusually preceded by dissociation of insulin oligomersinto monomers.38 Zinc ions exert their fibrillation-inhibitory effect by neutralizing negative charges inthe center of the insulin hexamer, whereby the hex-americ assembly is stabilized. Similar results havebeen reported when the addition of zinc was shownto prolong pulmonary insulin release from sodiumhyaluronate powder in beagle dogs up to 5 h comparedwith 3 h when no zinc was added.39 In another study,which demonstrates the effect of aggregation on bio-logical activity, prolonged insulin activity (manifestedby blood glucose reduction) was observed in nonagi-tated insulin gel as compared with the agitated in-sulin gel.40
The bioavailability of insulin in rats was cal-culated in terms of AUC values. When compared
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with subcutaneous insulin solution, insulin bioavail-ability was enhanced by 2.65- and 6.34-fold fromphase-sensitive delivery systems containing insulinalone and zinc–insulin, respectively. It was obviousupon evaluating In Vivo pharmacokinetic parametersthat delivery systems enhanced insulin bioavailabil-ity compared with subcutaneous aqueous insulin so-lution as denoted by the increase in AUC. The slowerabsorption rate leads to a higher percentage of insulindegradation by proteases in skin.41 In contrast, con-trolled delivery systems release small amount of in-sulin at a given time, which is immediately absorbed.In addition, polymeric systems offer additional stabi-lizing effect through limiting protein movement andprotection from surrounding environment.
Biocompatibility
In vitro cell culture studies have the advantage of rel-atively well-controlled assay conditions and are gen-erally accepted as an effective method for biocompat-ibility testing. In our study, phase-sensitive deliverysystems were extracted at 37◦C and 70◦C in PBS. Thelater group was carried out because polymers degradefaster at elevated temperatures, which simulate thelong-term effects of in situ depot degradation.42 Neu-tralized extracts of phase-sensitive delivery systemsin both conditions did not show cell growth inhibitionin the MTT assay in most of dilutions (Fig. 8). Therewas a slight reduction in cell proliferation betweenpolymer extracts in PBS at 1:1 and 1:2 dilutions. Athigher dilutions, there was no significant difference(p < 0.05) in cell viability between polymer extractsand PBS groups.
The decrease in cell viability in cells treated withphase-sensitive delivery system extracts at dilution1:1 and 1:2 is thought to be caused by the presence oftriacetin in the formulations. Yet, the “closed” incu-bation conditions do not simulate In Vivo conditionsin which degradation products are rapidly cleared.The exposure of limited number of cells to such highconcentrations of degradation products for such longperiods of time is not expected. Overall, the obtainedresults are good indications that the tested deliverysystems are biocompatible. However, in vitro biocom-patibility could vary fundamentally by altering celllines, exposure time, and growth media type. There-fore, it is important to examine the response of thetissue In Vivo to which the delivery system would beapplied.
An expected tissue reaction to polymeric implantswould be a short-lived inflammatory response to in-jury with minimal fibrosis resulting from wound heal-ing process. The normal tissue response to injectedimplants has been described by Shive and Anderson43
to occur in phases. First phase takes place withinthe first 2 weeks following injection and is charac-terized by the initiation and resolution of acute and
Figure 8. In vitro biocompatibility of phase-sensitivepolymer extracts. (a) Extracts prepared at 37◦C and (b) ex-tracts prepared at 70◦C (n = 8). PBS, phosphate bufferedsaline pH 7.4; PE, polymer extract; DMSO, dimethyl sul-foxide.
chronic inflammatory responses. This phase is distin-guished by the presence of neutrophiles, lymphocytes,and monocytes. After few days of injection, monocytesbecome the predominant cells. In the second phase,monocytes migrate to injury site and differentiate intomacrophages. Macrophages in turn may combine toform foreign body giant cells (FBGCs). The length oftime of the second phase depends on degradation rateand ability of macrophages, and if necessary FBGCs,to clear the tissue.
Histological evaluation of skin samples retrievedfrom injection sites of phase-sensitive delivery sys-tems showed a typical response to injury and normalwound healing process (Fig. 9). Acute inflammatoryresponse was obvious in day 1 samples, indicated bythe infiltration of neutrophiles and lymphocytes. Af-ter 1 week, there was a noticeable reduction in inflam-matory cells due to the resolution of acute response.Large numbers of macrophages were observed, whichindicate chronic inflammatory response to injury in-cidence. Less inflammatory cells were identified 1month after injection, signifying the subsiding of theinflammatory response. All signs of inflammatory
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2170 AL-TAHAMI, OAK, AND SINGH
Figure 9. Light micrographs of control skin sample (a)and samples following subcutaneous administration of poly-meric delivery system: (b) day 1, (c) day 7, (d) day 30, and(e) day 90.
responses diminished after 3 months, with connec-tive and muscular tissues, which were highly com-parable to control skin samples. According to the In-ternational Organization for Standardization (ISO)regulations,44 the final reactivity or local acceptance/rejection to implanted materials is evaluated histo-logically after approximately 3 months. At this time,the so-called steady state would have been reached.The In Vivo biocompatibility evaluation results in thisstudy are in agreement with in vitro MTT assay re-sults, and with ISO regulations, supporting the bio-compatible nature of the developed delivery systems.
In conclusion, the present work demonstrates thepotential of developing biocompatible and biodegrad-able delivery systems, which release insulin over ex-tended periods, after a single subcutaneous injectionin order to meet basal insulin requirements. The re-leased insulin was stable and conserved its biologicalactivity as demonstrated by the reduction in bloodglucose level.
ACKNOWLEDGMENTS
We acknowledge the financial support from NIH grantnumber HD056053 and the Fulbright Program forscholarship to Khaled Al-Tahami. We also acknowl-edge the assistance provided by Gitanjali Sharma inthe preparation of the manuscript.
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