injectable biodegradable hydrogels

Feature Article

Injectable Biodegradable Hydrogels

Minh Khanh Nguyen, Doo Sung Lee*

Injectable biodegradable copolymer hydrogels, which exhibit a sol–gel phase transition inresponse to external stimuli, such as temperature changes or both pH and temperature (pH/temperature) alterations, have found a number of uses in biomedical and pharmaceuticalapplications, such as drug delivery, cell growth, and tissue engineering. These hydrogels can beused in simple pharmaceutical formulations that can be prepared by mixing the hydrogel withdrugs, proteins, or cells. Such formulations are administered in a straightforward manner,through site-specific control of release behavior,and the hydrogels are compatible with biologicalsystems. This review will provide a summary ofrecent progress in biodegradable temperature-sensitive polymers including polyesters, polypho-sphazenes, polypeptides, and chitosan, and pH/temperature-sensitive polymers such as sulfa-methazine-, poly(b-amino ester)-, poly(aminourethane)-, and poly(amidoamine)-based polymers.The advantages of pH/temperature-sensitivepolymers over simple temperature-sensitive poly-mers are also discussed. A perspective on thefuture of injectable biodegradable hydrogels isoffered.

Introduction

Hydrogels are three-dimensional hydrophilic polymeric

networks that can absorb and retain a considerable amount

of water with maintenance of shape.[1,2] Injectable

biodegradable hydrogels have been widely used in

biomedical applications, such as drug/cell delivery and

tissue engineering, because of their highly hydrophilic

characteristics. Such hydrogels are of particular interest

because drugs, proteins, and cells can be easily incorporated

into polymer solutions prior to administration. Importantly,

D. S. Lee, M. K. NguyenDepartment of Polymer Science and Engineering, SungkyunkwanUniversity, Suwon, Gyeonggi 440-476, KoreaFax: þ82-31-292-8790; E-mail: [email protected]

Macromol. Biosci. 2010, 10, 563–579

� 2010 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

no surgical procedures are required for the insertion of gels

into the body; the gels are administered by simple

injection.[3–5]

Injectable hydrogels can be formed in situ by either

chemical or physical crosslinking methods. Chemically

crosslinked hydrogels, prepared through photopolymeriza-

tion,[6] disulfide bond formation,[7] or reaction between

thiols and acrylate or sulfones,[7] undergo significant

volume changes during the phase transition. In contrast

to chemical hydrogels, physically crosslinked hydrogels,

formed by the self-assembly of polymers in response to

environmental stimuli (for example, temperature, pH, or

both) display sol–gel transitions without marked volume

changes. The sol–gel transition systems are relatively low-

viscosity aqueous solutions (sol state) prior to injection, but

rapidly convert into a gel under physiological conditions

DOI: 10.1002/mabi.200900402 563

M. K. Nguyen, D. S. Lee

Minh Khanh Nguyen graduated from theDepartment of Chemical Engineering fromHoChiMinh City University of Technology (Viet-Nam) in 2003 and is currently pursuing his Ph.D.under the guidance of Professor Doo Sung Lee atSungkyunkwan University (Korea). His mainresearch is focused on the development of func-tionalized and biodegradable injectable poly-meric hydrogels for controlled drug andprotein delivery.

Doo Sung Lee studied chemical engineering atthe Seoul National University. He completed hisM.Sc. and Ph.D. at Korea Advanced Institute ofScience and Technology in 1984 working in thefield of interpenetrating polymer networks. Hejoined Sungkyunkwan University as AssistantProfessor in the Department of Textile Engineer-ing. He was one of the founders of the Depart-ment of Polymer Science and Engineering andwas appointed Professor of this department in1993. He served as a Dean of the College ofEngineering at Sungkyunkwan University in2005–2007. He started to study biomaterials in

564

post-injection.[1–3] Physical hydrogels have several advan-

tages over chemical hydrogels, because they do not require

photo irradiation, use of organic solvents or crosslinking

agents, and do not release heat during polymerization at the

gelation site, which may denature incorporated proteins

and damage embedded cells and surrounding tissues. Thus,

the physical systems have recently attracted increased

attention. For use in drug/cell delivery and tissue engineer-

ing, hydrogels should be low-viscosity solutions (free-

flowing) prior to subcutaneous injection, and should

rapidly gel in the human body, where ultimate degradation

of the hydrogels is desired.

This article provides insights into recent advances in

synthesis and biomedical applications of injectable, biode-

gradable, polymeric hydrogels that exhibit sol–gel transi-

tions in response to temperature and pH/temperature

changes. The discussion covers poly(ethylene glycol) (PEG)/

polyester block copolymers, polyphosphazenes, polypep-

tides, chitosan, polymers based on sulfamethazine, poly(b-

amino ester), poly(amino urethane), poly(amidoamine),

and others.

1995 when he was in the Department of Phar-maceutics and Pharmaceutical Chemistry at theUniversity of Utah as a visiting Professor. Heserved as an Editor-in-chief of Polymer Scienceand Technology (2001) and on the Editorial boardof Macromolecular Research for four years(2002–2005). He was elected as a member ofKorean Academy of Engineering in 2007. He iscurrently a vice president and an editorial boardmember of Biomaterials Research of the KoreanSociety of Biomaterals (2006–present). He hasauthored and co-authored about 300 papers,including about 120 in peer-reviewed journals,and seven book chapters, and filed 26 patents.His main research interest is functionalized andbiodegradable injectable hydrogels and micellesfor the controlled drug and protein delivery andmolecular imaging.

Injectable Biodegradable Block CopolymerHydrogels

Thermosensitive Block Copolymer Hydrogels

Hydrogels that are sensitive to temperature are useful for

both in vitro and in vivo applications, because temperature

control is generally easy. Temperature-sensitive hydrogels

undergo a sol–gel phase transition when the temperature is

increased from room temperature to physiological tem-

peratures.

Poly(ethylene glycol) (PEG)/Polyester

Copolymers of hydrophilic biocompatible PEG with biode-

gradable biocompatible aliphatic polyesters, for example

polylactide (PLA), polyglycolide (PGA), poly(e-caprolactone)

(PCL), or poly[(R)-3-hydroxybutyrate] (PHB), have received

increasing attention as promising biomaterials.

The first reported biodegradable thermosensitive

hydrogel was PEG-poly(L-lactide)-PEG (PEG-PLLA-PEG)

(Scheme 1).[9] PEG-PLLA-PEG was synthesized by the ring-

opening polymerization (ROP) of L-lactide using mono-

methoxy PEG (MPEG) of molecular weight (MW) 5 000 as

macroinitiator, to form MPEG-PLLA diblock copolymers.

Triblock copolymers (PEG-PLLA-PEG), with PLLA blocks in

Scheme 1. Structure of PEG-PLLA-PEG.



the 2 000–5 000 Da molecular weight range were next

produced by coupling the resulting diblock copolymers

using hexamethylene diisocyanate (HMDI). The concen-

trated copolymers (10–30 wt.-%) dissolved in water

demonstrated a gel-to-sol transition with increases in

temperature. The gel-to-sol transition was precisely con-

trolled by the biodegradable PLLA block length when a PEG

block was fixed at both ends. The gel-to-sol properties of the

PEG/polyester diblock and triblock copolymers in water

depended on the hydrophobic/hydrophilic balance, block

length, hydrophobicity, and stereoregularity of the hydro-

phobic block.[10,11]

PEG-(D,L-lactide)-PEG (PEG-PLA-PEG) triblock copolymers

(MW of MPEG¼ 2 000–5 000) were recently produced by

DOI: 10.1002/mabi.200900402


Figure 1. Phase diagram of PEG-PLGA-PEG triblock copolymeraqueous solutions. Sol (flow) to gel (no flow) transition tempera-ture was measured by the test tube inverting method increasing2 8C/step. Reprinted with permission from ref.[16]. Copyright 1999Elsevier.

coupling MPEG-PLA diblock copolymers using adipoyl

chloride.[12] Rheological measurements showed a gel-to-

sol transition in concentrated polymer solutions, with an

increase in temperature. Gelation at lower temperatures

was attributed to hydrogen bonding between PEG blocks,

and the hydrogen bonds were broken at elevated tempera-

tures, leading to the gel-to-sol transition. A series of

biodegradable star-shaped PLLA-PEG block copolymers

were obtained by the pairing of two components: star

PLLA and monocarboxy-MPEG, using dicyclohexylcarbodii-

mide (DCC).[13,14] The gel-to-sol transitions, induced by

increases in temperature, were observed at concentrations

beyond the critical gelation concentration (CGC). With the

same PEG block length, use of longer PLLA blocks led to a

decrease in the CGC, thus widening the gelation window.

Gelation was attributed to micellar packing and the

breaking of micellar packing structures caused by the

partial hydration of the PEG block, which resulted in the gel-

to-sol transition at higher temperatures.

However, the abovementioned gels underwent gel-to-sol

transitions, reducing their suitability for encapsulation of

some drugs or proteins. In addition, injection at tempera-

tures that are elevated with respect to body temperature

is uncomfortable for patients. PEG-poly(D,L-lactide-co-

glycolide)-PEG (PEG-PLGA-PEG) triblock copolymers with

short PEG blocks (MW� 750) were soluble in water at low

temperatures and converted into a gel at elevated

temperatures (Scheme 2).[15] The triblock copolymers were

shown to be biocompatible with blood.[16–18] The gelation

window spanned the physiological temperature range

(Figure 1). 13C NMR and dynamic light scattering (DLS)

studies revealed that the gelation of PEG-PLGA-PEG was

enhanced by micellar growth and close packing of micelles

(Figure 2). The upper gel-to-sol transition at higher

temperatures was driven by the breakage of the micellar

structure caused by the partial dehydration of PEG and

PLGA blocks. The sol–gel transition could be adjusted by

variation in PLGA and PEG block lengths, PLGA composition,

and the use of particular additives. In other studies,[19,20] the

gelation mechanism of PEG-PLGA-PEG (550-2810-550) was

investigated by rheology, DLS, differential scanning calori-

metry (DSC), and small-angle neutron scattering (SANS).

The results indicated that the macroscopic liquid–liquid

phase separation induced gelation of the triblock copoly-

mer. A transparent gel was formed in situ after injection of a

33 wt.-% PEG-PLGA-PEG (550-2810-550) aqueous solution

into rats, and gel integrity persisted for one month.[21]

Interestingly, thermosensitive hydrogels based on PLGA-

PEG-PLGA (BAB-type) (Scheme 3) showed a sol-to-gel

Scheme 2. Structure of PEG-PLGA-PEG.



transition similar to that of PEG-PLGA-PEG (ABA-type)

copolymer hydrogels.[22–24] However, the synthetic proce-

dure for BAB-type hydrogels was simpler than that for the

ABA-type, using HMDI as a coupling agent. The gelation

mechanism for BAB-type hydrogels was different from that

of ABA-type hydrogels because of the presence of two PLGA

end blocks. The former formed micelles with intermicellar

bridges, whereas the latter formed regular micelles with a

PLGA block core and a PEG block shell, in aqueous solution.

At temperatures below the critical gelation temperature

(CGT), some bridging micelles of PLGA-PEG-PLGA copoly-

mers formed, but they were not stable because of the low

hydrophobicity of PLGA. With increasing temperatures to

the CGT, a bridged micelle network was formed because of

an increase in the hydrophobicity of the PLGA segment,

leading to gelation (Figure 3). In vitro and in vivo

degradation of the PLGA-PEG-PLGA (1500-1000-1500)

copolymer (ReGel) was studied.[23] A 23 wt.-% ReGel was

found to degrade completely in vitro after 6–8 weeks at

37 8C, whereas no hydrogels were observed at the end of the

fourth week after subcutaneous injection of a 23 wt.-%

ReGel solution into rats.

To overcome the molecular weight constraints of the

PLGA-PEG-PLGA and PEG-PLGA-PEG triblock copolymers,

and to control the time of persistence of gels, graft

copolymers of PEG-g-PLGA and PLGA-g-PEG were investi-

gated (Scheme 4).[25–27] The former copolymer was

synthesized by the ROP of D,L-lactide (LA) and glycolide

(GA) using a hydroxy-pendant PEG as a macroinitiator. The

PEG-g-PLGA copolymer in water showed a sol–gel transi-

tion at concentrations above 16 wt.-%. Gel integrity

persisted for one week under physiological conditions.

The latter copolymer was prepared by the one-step ROP of

LA, GA, and epoxy-terminated PEG.[25,26] Aqueous solutions

www.mbs-journal.de 565


Figure 2. Schematic diagram of the sol-gel transition of PEG-PLGA-PEG aqueous solutionin response to temperature.

Scheme 5. PCL-PEG-PCL.

Scheme 3. Structure of PLGA-PEG-PLGA.

Figure 3. Schematic diagram of the sol–gel transition of PLGA-PEG-PLGA aqueous solution in response to temperature.

Scheme 4. Chemical structures of a) PEG-g-PLGA and b) PLGA-g-PEG.

566

of PLGA-g-PEG copolymers also exhibited a sol-to-gel

transition upon heating. After injection of the copolymer

solution (29 wt.-%) into rats, the gel persisted for more than

two months, a significant improvement over the one week

persistence time observed for the PEG-g-PLGA copolymer

hydrogel. The gelation mechanism of the PLGA-g-PEG

copolymer was examined by 13C NMR spectroscopy, SANS,

rheology, and infrared (IR) spectroscopy.[28] Partial dehy-

dration of PEG caused micellar aggregation, leading to

gelation, and significant dehydration of PEG led to

macroscopic separation at the gel-to-sol transition



temperature. The sol-to-gel transition

temperature was tuned from 15 to

45 8C by tailoring the polymer composi-

tion, the number of PEG grafts present,

and the PEG molecular weight. By mixing

these copolymers at different composi-

tion ratios, the sol-to-gel transition tem-

perature could be adjusted, and the gel

duration could be varied from one week

to three months. Both copolymers

yielded rather soft gels with storage

moduli (G0 values) of 100 Pa.[26]

Poly(e-caprolactone) (PCL) is a hydro-

phobic crystalline polymer that is both

biodegradable and biocompatible. PCL copolymers have a

powdery morphology, making them easier to handle than

are PLGA and PLLA copolymers, which have a sticky paste

morphology. Aqueous solutions of MPEG-PCL diblock

copolymers (MPEG MW� 2 000) underwent a gel-to-sol

transition on temperature increase. The gelation window

was strongly influenced by PEG and PCL block length. In

vivo gelation of MPEG-PCL (MW 2 000–2 300) was also

studied. When a 23 wt.-% copolymer solution at 42 8C was

injected into a rat, the gel formed immediately and

persisted for one month with only a small amount of

inflammation at the injection site.[29] MPEG-PCL diblock

copolymers with a lower molecular weight MPEG

(MW¼ 750) were subsequently reported.[30] Interestingly,

synthesized diblock copolymers with PCL block lengths of

1 400–3 000 were soluble in water and underwent a sol-to-

gel-to-sol transition as a result of temperature changes. The

sol–gel phase diagram depended on PCL block length.

Recently, PEG-PCL-PEG and PCL-PEG-PCL (Scheme 5) triblock

copolymers have been introduced, and aqueous solutions

thereof showed a clear sol-to-gel-to-turbid sol transition

with an increase in temperature.[31,32] DLS and 13C NMR

studies indicated that the clear sol-to-gel transition arose

from micellar aggregation, whereas the gel-to-turbid sol

transition was driven by the breakage of the core–shell

structure. Because of differences in structural topology, PCL-

PEG-PCL had a higher G0 (10 000 Pa) (Figure 4) than did

PEG-PCL-PEG (100 Pa) (Figure 5). However, because of

crystallization of the PCL block, a clear aqueous solution

of PCL-PEG-PCL (20 wt.-%) turned turbid within 1 h at 20 8C,

which may be problematic for injectability. The crystal-

lization problem was, however, solved by using poly-

(caprolactone-co-trimethylene carbonate)-PEG-poly(capro-

lactone-co-trimethylene carbonate) (PCTC-PEG-PCTC).[33] A

DOI: 10.1002/mabi.200900402


Figure 4. Dynamic mechanical analysis of PCL-PEG-PCL triblockcopolymer aqueous solutions (20wt.-%) as a function of tempera-ture. The thermogram was obtained with a heating rate of0.2 8C �min�1. Reprinted with permission from ref.[31]. Copyright2005 American Chemical Society.

Figure 5. Dynamic mechanical analysis of the PEG-PCL-PEG tri-block copolymer aqueous solutions as a function of temperatureand concentration. Reprinted with permission from ref.[30]. Copy-right 2005 American Chemical Society.

PCTC-PEG-PCTC triblock copolymer in water exhibited a sol-

to-gel-to-syneresis transition upon heating. However, a

very soft gel with a G0 of 1 Pa was obtained. This gel was

stable with respect to hydrolysis in phosphate buffers at

37 8C for 50 d, but degraded markedly in rats. In addition, the

crystallizability of PCL decreased as the molecular weight of

PCL rose. Therefore, a multiblock copolymer of (PCL-PEG-

PCL)n was prepared by coupling PCL-PEG-PCL triblock

copolymers (1 000-1 000-1 000), using terephthaloyl chlor-

ide. A 20 wt.-% aqueous solution of the multiblock

copolymer underwent a sol-to-gel-to-sol transition that

did not turn turbid at room temperature, indicating that the

material may be convenient for practical applications, such

as drug formulation and injection. The multiblock showed a

lower G0 (100 Pa) than did the triblock (10 000 Pa). The

gelation mechanism was attributed to crystallization of the

multiblock copolymer.[34]

The multiblock topology affected the gelation behavior

of other copolymer hydrogels. A series of PEG/PLLA

alternating multiblock copolymers was synthesized by

coupling PEG (MW¼ 600) to PLLA (MW¼ 1 100–1 500) using

succinic anhydride.[35] The multiblock PEG/PLLA aqueous



solution underwent a sol-to-gel-to-sol transition with

increasing temperature. The transition temperature and

gel modulus could be controlled by varying the PLLA block

length, PEG molecular weight, and PEG/PLLA ratio. The

gelation mechanism was considered to be micelle aggrega-

tion. Stereochemistry also affected gelation.[36] The PEG/

poly(D,L-lactide) (PDLLA) and PEG/PLLA multiblock copoly-

mers with identical block lengths and total molecular

weight were prepared. Relative to amorphous PEG/PDLLA,

the stereoregular PEG/PLLA multiblock copolymer had a

reduced CGC, a lower sol-to-gel transition temperature, a

broader gel area, and a larger maximal gel modulus.13C NMR, X-ray diffraction (XRD), and UV-visible spectro-

scopy studies indicated that the different gelation proper-

ties could be attributed to the slower dynamic molecular

motion of the methyl groups in PLLA. The isotactic

arrangement in PLLA induced strong aggregation of the

PEG/PLLA multiblock. In contrast with the sol-to-gel

transition behavior of the PCL-PEG-PCL (1 000-1 000-1 000)

multiblock copolymer,[34] the PCL-PEG-PCL multiblock

copolymer, with a higher molecular weight of both PEG

and PCL, exhibited a gel-to-sol transition upon heating.[37]

In addition, multiblock copolymers of PEG-sebacate (PEG-

SA) were synthesized by simple condensation polymeriza-

tion.[38] A soft gel formed when a 25 wt.-% aqueous solution

of the PEG-SA multiblock was heated to 37 8C, and gel

integrity persisted for more than three weeks in phosphate

buffer, pH 7.4, at 37 8C.

In addition, it was found that a stereocomplex of the

enantiomeric triblock copolymers (10 wt.-%), PLLA-PEG-

PLLA (1 300-4 600-1 300), and poly(D-lactide)-PEG-poly-

(D-lactide) (PDLA-PEG-PDLA) (1 100-4 600-1 100), in water

could induce temperature-dependent gelation, although

the individual enantiomeric copolymers did not show

thermal gelation. Stereocomplex formation during gelation

was confirmed by wide-angle X-ray scattering (WAXS).[39] A

mixture of PEG-PLLA-PEG (2 000-2 000-2 000) and PEG-

PDLA-PEG (2 000-2 000-2 000) (35 wt.-%) exhibited a gel-

to-sol transition with an increase in temperature. WAXS

results showed that gelation was not attributable to

complexation of PLLA and PDLA, but rather to interdigita-

tion of the helical PEG chains induced by the complemen-

tary arrangement of PDLA and PLLA helices.[40]

Enantiomeric PEG12500-(PLA)2 and PEG21800-(PLA)8 aqu-

eous solutions showed gel-to-sol transitions with increas-

ing temperatures. Although neither enantiomeric polymer

solution formed a gel, a stereocomplex with a PLA block

could form a gel at the same concentration. Rheology

studies revealed that PEG-(PLA)8 showed a higher gel

modulus than PEG-(PLA)2, because of the higher cross-

linking density of the former. The PLA block length, PEG

content, polymer topology, and polymer concentration

influenced the gel–sol transition, gel modulus, and kinetics

of gelation.[41,42] Subsequently, the PEG-PLLA and PEG-PDLA



568

multiblock copolymers were developed by coupling the

corresponding triblock copolymers using diisocyanatobu-

tane.[43] The multiblock copolymers showed a much lower

CGC, faster gelation, and a higher gel storage modulus,

compared with the parent triblock copolymers, because of

an increase in the crosslinking density of the multiblock

copolymers.

The end groups of the thermosensitive copolymers also

influenced gelation.[44,45] A series of PLGA-PEG-PLGA tri-

block copolymers with several different end caps (hydroxy,

acetyl, propionyl, and butanoyl groups) was synthesized

and characterized. Triblock copolymers containing acetate

and propionate groups exhibited a sol-to-gel transition as a

function of temperature, whereas the copolymer contain-

ing the butyrate group precipitated in water. An increase in

the hydrophobicity of the copolymer lowered the transition

temperature, and the end cap caused a significant change

in the position of the gelation window. Cholesterol end-

capped star PEG-PLLA copolymers (above 3 wt.-%) in water

also exhibited thermal gelation, but PEG-PLLA itself did

not.[46] Gelation was induced by the strong hydrophobic

association of the cholesterol group.

Poly[(R)-3-hydroxybutyrate] (PHB) is a natural biode-

gradable polyester produced by bacteria. The crystallinity

and hydrophobicity were higher than those of other

synthetic polymers, such as PLA and PCL. A thermosensitive

and amphiphilic poly(ether ester urethane) multiblock

copolymer consisting of PHB, PEG, and poly(propylene

glycol) (PPG) (PEG/PPG/PHB) was synthesized.[47,48] This

copolymer, in aqueous solutions and at very low concen-

trations (2–5 wt.-%). showed a sol-to-gel transition as a

function of temperature change. The gelation of the

copolymer solution was associated with micellar packing.

Polyphosphazenes

Biodegradable polyphosphazenes, consisting of a hydro-

philic PEG block and hydrophobic amino acids or a peptide

block, such as L-isoleucine ethyl ester (IleOEt), D,L-leucine

ethyl ester (LeuOEt), L-valine ethyl ester (ValOEt), or di-, tri-,

and oligo-peptides in the side groups, were synthesized

(Scheme 6).[49–51] Aqueous solutions (10 wt.-%) of polypho-

sphazenes with MPEG350 and IleOEt exhibited a sol-to-gel

transition as a function of temperature.[49] The maximal

viscosity was 30 Pa � s at 37 8C. The gelation properties were

Scheme 6. Structure of polyphosphazene.



adjusted by varying the composition of the substituents,

MPEG molecular weight, and concentration. In a subse-

quent study, polyphosphazenes with oligopeptides (tri- or

tetra-peptides) and MPEG 350 as side groups were also

found to exhibit a phase transition.[50] The gelation

properties of the thermogelling polyphosphazenes

depended on the structure of the oligopeptide and the

hydrophobic side groups. Such gels also showed higher gel

strength compared with former gels.[49] Intermolecular

association of hydrophobic oligopeptides was responsible

for thermally induced gelation of the copolymer. The sol–

gel transition and gel strength of polyphosphazenes were

also modulated by the blending of hard and soft polymers.

By mixing the two polymers at Tmax (the temperature at

which the viscosity is maximal) values of 31 and 42 8C, with

blend ratios of 2:2, the Tmax of the blended polymers was

observed to be 35–41 8C.[52] The degradation rate could be

controlled by the content of incorporated depsipeptides.

Polyphosphazenes, incorporated with the depsipeptides,

degraded faster than those in gels without

depsipeptides, because the hydrolysis of depsipeptides

produced carboxylic acid, which triggered degradation of

hydrophobic amino acid groups.[51]

Polypeptides

Polypeptides are important biomaterials offering favorable

characteristics, such as biocompatibility and biodegrad-

ability. Polypeptides can interconvert among a variety of

conformations, such as a-helix, b-sheet, and random coil,

and building blocks with hydrophobic, hydrophilic, ionic,

and non-ionic characteristics can be synthesized. An

artificial triblock protein with short leucine-zipper end

blocks flanking a water-soluble polyelectrolyte domain

underwent reversible gelation in response to changes in pH

and temperature.[53] Gelation of the triblock protein was

driven by formation of coiled-coil aggregates of the

terminal leucine-zipper domains, and the gel changed to

a viscous solution when coiled-coil aggregates dissociated

with increasing pH and temperature. The thermally

induced hybrid hydrogels were prepared by combination

of water-soluble synthetic polymers and engineered

proteins.[54] The proteins were used as crosslinkers for

synthetic polymers. Changes in protein conformation as a

result of temperature changes triggered the formation of

hybrid hydrogels. Diblock copolypeptide amphiphiles,

consisting of charged and hydrophobic blocks, were

synthesized.[55] Aqueous solutions could form thermally

stable gels (up to 90 8C), but the gels rapidly broke down

under an applied stress. Gelation was believed to be

promoted by association of the hydrophobic domains,

which was triggered by the ordered packing ofa-helical and

b-strand segments, whereas rapid recovery after stress was

attributable to the nature of the physical gelation process

and the low molecular weight of the copolypeptides. A

DOI: 10.1002/mabi.200900402

surplus2

Highlight


Scheme 8. Chitosan structure.

10 wt.-% b-lactoglobulin aqueous solution showed a sol-to-

gel transition at 85 8C.[56] b-lactoglobulin is a component of

milk whey, with two disulfide bonds and one free

sulfhydryl group. Gelation was attributed to the formation

of hydrophobically linked aggregates, followed by forma-

tion of disulfide-bonded aggregates. A synthetic polypep-

tide of poly(ferrocenylsilane)-poly(g-benzyl-L-glutamate)

(PFS-PBLG) was soluble in hot toluene, but formed a

transparent gel at room temperature.[57] The gelation

process arose from formation of an a-helix and random-

coil structure in the diblock copolymer in toluene.

VPGVG (V¼ valine, P¼proline, and G¼ glycine) is a

prominent amino acid sequence in elastin. A hexahistidine

metal-binding motif was incorporated into an elastin-like

polypeptide.[58] A 6 wt.-% aqueous solution of the polymer

was a clear at 4 8C, but changed to a gel at 25 8C. This

material may be useful in heavy metal removal applica-

tions. Copolymers that contained a collagen peptide and an

elastin peptide formed a thermally induced gel in water.[59]

Polypeptides with 82–86 mol-% VPGVG composition

showed a sol-to-gel transition when the temperature was

increased. The collagen acted as a hydrate unit and the

elastin peptide acted as a thermosensitive crosslinking

point. A de novo-designed peptide showed a thermosensi-

tive gelation transition.[60] A 2 wt.-% aqueous solution of

the MAX3 peptide was a rigid gel at 75 8C withG0 of 1 100 Pa.

A transition from a random coil to a b-hairpin produced a

hydrogel network as the temperature increased. A sol-to-gel

transition was also evident in aqueous solutions (above

3 wt.-%) of amphiphilic poly(N-substituted a/b-aspara-

gines) (Scheme 7).[61] The phase diagram was strongly

influenced by hydrophilic blocks (amino alcohols) and

polymer concentration. Recently, poly(alanine)-poloxamer-

poly(alanine) (PA-PLX-PA) was synthesized as a thermo-

sensitive hydrogel.[62] Aqueous solutions of PA-PLX-PA

underwent a sol-to-gel transition as the temperature

increased. The sol-to-gel transition temperature was

influenced by the molecular weight of each block and by

the composition of PA. Based on FT-IR, DLS, 13C NMR, circular

dichroism (CD), transmission electron microscopy (TEM),

and fluorescence spectroscopy studies, the PA transition

from random coil to b-sheet and the decrease in molecular

motion of PLX resulted in gelation. The hydrogels were

stable in phosphate buffer but degraded quickly in the

Scheme 7. Structure of poly(N-substituted a/b-asparagine).



presence of enzymes. A new thermogelling poly(N-vinyl

pyrrolidone)-PA (PVP-PA) has been reported.[63] Aqueous

solutions of the polymers showed sol-to-gel transitions. Gel

formation was attributed to hydrophobic association and

formation of a b-sheet structure.

Chitosan

Chitosan, a polysaccharide derived from the partial

deacetylation of chitin from crustacean shells, has been

widely used as a biomaterial because of biodegradability,

biocompatibility, non-toxicity, and bioadhesive properties

(Scheme 8). Chitosan was approved by the US Food and Drug

Administration and has been used in drug delivery, tissue

engineering, and cosmetics.[64,65] An injectable thermogel-

ling hydrogel was prepared by the combination of chitosan

andb-glycerol phosphate (C/GP).[66] Chitosan was dissolved

in hydrochloric acid, and a GP solution was then slowly

added to obtain a clear solution. At pH 7.15, the C/GP

aqueous solution remained in a clear liquid state, but gelled

rapidly in the vicinity of 37 8C when heated. The gelation

temperature increased as the degree of deacetylation

decreased, but was not affected by the molecular weight

of the chitosan. Gelation was driven by hydrophobic

association of the neutral chitosan molecules, promoted

by the influence of GP on water at elevated temperatures.

When subcutaneously injected into rats, the C/GP solution

rapidly gelled. PEG-g-chitosan aqueous solutions exhibited

thermal gelation.[67] A 3 wt.-% solution of chitosan grafted

onto 55 wt.-% PEG showed an increase in viscosity at 37 8C,

from < 1 Pa � s to > 6 Pa � s within 1 250 s, whereas the

chitosan solution alone did not show any change in

viscosity, even after 3 500 s. A network gel resulted from

the association of chitosan molecules and reduction in the

mobility of the PEG segments at high temperatures.

Pluronic was grafted onto chitosan (C-g-P), and the resulting

C-g-P in aqueous solution displayed a sol-to-gel transition

as the temperature increased.[68] The gelation temperature

was controlled by chitosan content, and gelation did not

occur when the chitosan content was >17 wt.-%. More

recently, hydrophobic N-palmitoyl moieties were grafted

onto chitosan (NPCS) to produce a pH-triggered hydrogel

within the pH range of 6.5–7.0.[69] The G0 of the NPCS

aqueous solution at pH 6.5 was about 100 Pa (Figure 6) and

was strongly dependent on the shear rate, indicating that



Figure 6. Dynamic temperature sweeps of aqueous NPCS(1% w/v, pH 6.5) at an oscillatory strain amplitude of 1% and afrequency of 0.1 Hz. Reprinted with permission from ref.[68].Copyright 2009 Elsevier.

570

the NPCS aqueous solution might be able to squeeze

through the needle during injection. The G0 of the hydrogel

at pH 7.0 was higher than that at pH 6.5. This material was

non-toxic in vitro, but a large degree of inflammation was

observed at the interface between the tissue and the

hydrogel after two weeks of implantation. No chronic

inflammation was observed after six weeks of implanta-

tion. A balance between charge repulsion and hydrophobic

interactions in NPCS aqueous solutions in the pH range 6.5–

7.0 was involved in the gelation process.

Other Thermosensitive Block Copolymers

Poly(trimethylene carbonate) (PTMC) is biodegradable,

biocompatible, and has soft mechanical properties. A

PEG-PTMC diblock copolymer was synthesized by the

ROP of trimethylene carbonate (TMC) onto MPEG, using

stannous octoate as a catalyst.[70] The diblock copolymer

solution (�25 wt.-%) underwent a sol-to-gel transition on

an increase in temperature. The phase diagram was

mapped by varying the concentration, molecular weight,

and composition of the diblock copolymer. On the basis of13C NMR, DLS, and TEM studies, micellar aggregation,

through dehydration of PEG, was attributed to the gelation

process. The hydrogel was stable in vitro for up to 90 days,

but a 15% weight loss occurred over 20 days in vivo. The

degradation in vivo, which is different from the degrada-

tion in vitro, may be due to the influence of body fluid. The

degradation of PTMC in vivo produced alcohol and carbon

dioxide, which did not reduce the pH at the interface

between the hydrogel and the tissue.[71] ABA-type block

copolymers consisting of poly(propylene fumarate) (PPF)

and MPEG were synthesized by a simple transesterification

method.[72] The triblock copolymer with MPEG molecular

weights of 570 and 800 in aqueous solution exhibited a

thermosensitive gelation process in the concentration



range of 5–25 wt.-%. The sol–gel transition was influenced

by salt concentration and MPEG molecular weight. Highly

unsaturated double bonds in PPF could form in situ

crosslinks. Poly(propylene phosphate) (PPP) has been used

in biomaterial applications, such as drug delivery, tissue

engineering, and gene delivery, because of its biodegrad-

ability and biocompatibility.[73] PPP aqueous solutions did

not exhibit a thermally induced gelation transition, but

underwent a sol-to-gel transition in the presence of calcium

ions. Polyacetal grafted with MPEG was prepared, to make a

thermogelling hydrogel.[74] When a 15% graft ratio of MPEG

and a 5% graft ratio of poly(orthoester) were introduced

onto the polyacetal backbone, the resulting polymer in

aqueous solutions (25 wt.-%) underwent a sol-to-gel transi-

tion at 34 8C. PEG-poly(ethyl-2-cyanoacrylate) (PEG-PEC)

was synthesized by addition polymerization.[75] A PEG-PEC

aqueous solution (750-450) showed a unique closed-loop

phase transition in the concentration range of 4–15 wt.-%.

In contrast with the systems described above, such as PCL

and PLGA, the sol-to-gel transition temperature of PEG-PEC

increased when the polymer concentration increased.

Closed-loop gelation resulted from a balance between

aggregation and stabilization of micelles as a function of

gelation temperature and concentration.

pH/Temperature-Sensitive Block CopolymerHydrogels

The thermosensitive block copolymer hydrogels have

potential applications as biomaterials. However, they

suffer from some limitations that restrict the range of

applications in which they may be utilized. First, when a

thermosensitive polymer solution is injected into the body

using a syringe, the increase in temperature to the

physiological temperature (37 8C) during injection causes

gelation inside the needle, creating a blockage. This makes it

difficult to inject thermosensitive polymer solutions into

the body. Second, a lack of functional groups limits

applications of these materials with respect to delivery of

ionic peptides/proteins. Third, it takes a long time to

dissolve the thermosensitive polymers in water, thus the

polymer should be stored prior to use. The biodegradable

polymer could be degraded during storage and circulation

for commercial use. Therefore, the reconstitution problem

of the polymer solution is of concern. Fourth, the degrada-

tion of polyester generates acidic products sometimes that

change the local pH. The resulting low pH damages

incorporated proteins or cells. Thus, it is important to

maintain neutral pH during the degradation.

pH/temperature-sensitive copolymer hydrogels were

prepared by combining a pH-sensitive moiety with a

temperature-sensitive block to solve the abovementioned

drawbacks. Acidic sulfamethazine oligomers (OSMs) were

DOI: 10.1002/mabi.200900402


Scheme 9. Structure of OSM-PCLA-PEG-PCLA-OSM.

Figure 7. Phase diagram of block copolymers in buffer solution.Mn of PEG¼ 1 750; concentration, 15wt.-%; PEG/PCLA weightratio, 1/1.89 (&), 1/2.08 (~). a) PCLA-PEG-PCLA solutionb) OSM-PCLA-PEG-PCLA-OSM solution. A) pH 7.4, 37 8C; B) pH8.0, 37 8C; C) pH 7.4, 15 8C; D) pH 8.0, 15 8C. Reprinted with per-mission from ref.[75]. Copyright 2005 American Chemical Society.

coupled with thermosensitive poly(e-CL-co-LA)-PEG-poly-

(e-CL-co-LA) triblock copolymers to produce pH/tempera-

ture-sensitive hydrogels (OSM-PCLA-PEG-PCLA-OSM)

(Scheme 9).[76] These copolymer hydrogels were synthe-

sized in two steps: First, a carboxylic acid-terminated OSM

was obtained by conventional radical polymerization in the

presence of a chain transfer agent (3-mercaptopropionic

acid), and a PCLA-PEG-PCLA triblock copolymer was

produced by the ROP of CL and LA using PEG as a

macroinitiator. Second, the carboxylic group in OSM was

coupled to the hydroxy groups at both ends of PCLA-PEG-

PCLA using 4-(dimethylamino) pyridine (DMAP) as a

catalyst. The parent PCLA-PEG-PCLA triblock copolymer

aqueous solutions (15 wt.-%) showed a sol-to-gel transition

in response to changes in temperature but not in pH

(Figure 7a). In contrast, the 15 wt.-% OSM-PCLA-PEG-PCLA-

OSM solution exhibited a sol-to-gel transition as a function

of both pH and temperature (Figure 7b). The gel window

became wider with increasing pH and/or PCLA/PEG ratio.

The sol–gel transition could be controlled by the polymer

concentration, hydrophobic/hydrophilic balance, PEG block

length, and OSM molecular weight.[77] An association of

bridged micelles was suggested as a mechanism for

gelation of the pentablock copolymer. A schematic gelation

mechanism is illustrated in Figure 8. At pH 8.0 and in the

temperature range of 10–70 8C, the polymer solution

existed as a sol state because the OSM was ionized. At

pH 7.4 and 15 8C, the OSM deionized and became more

hydrophobic, but the pentablock copolymer still exhibited a

sol state because of weak interactions with the hydrophilic

PCLA block at low temperatures. In contrast, at pH 7.4 and

37 8C, the PCLA blocks became hydrophobic, inducing a

strong hydrophobic interaction between PCLA-OSM blocks

and leading to a micellar interconnecting gelation process.

At pH 8.0, the polymer solution did not form a gel in the

temperature range of 10–70 8C, suggesting that the solution

was easily injectable using a syringe. A strong gel formed

quickly when the polymer solution was injected into a

pH 7.4 phosphate buffered saline (PBS) solution, whereas

dispersion of the polymer in a pH 8.0 PBS solution was

observed. The pentablock copolymer hydrogel maintained

its integrity for more than two weeks in pH 7.4 PBS at 37 8Cand showed a slower degradation than did the parent PCLA-

PEG-PCLA polymer.[78] After one month, the molecular



weight of the pentablock copoly-

mer decreased from 6 550 to 4 830.

The pH drop as a result of degrada-

tion of the parent PCLA-PEG-PCLA

triblock copolymer was significant,

from pH 7.4 to 2.2, whereas the pH

drop resulting from OSM-PCLA-

PEG-PCLA-OSM degradation was

from only pH 7.4 to 5.5 after one

month. The buffering effect of OSM

moieties minimized the effects of acidic degradation

products. A hydrogel formed rapidly after injection of the

pentablock copolymer solution (20 wt.-%, pH 8.0) into rats.

Good cytotoxicity against HeLa cells was observed in vitro

for concentrations up to 10 mg �mL�1 of the pentablock

copolymer. A histology study revealed that acute inflam-

mation was found during the first two weeks, but decreased

notably after six weeks. Well-defined OSM-PCLA-PEG-PCLA-

OSM pentablock copolymers were synthesized by atom

transfer radial polymerization (ATRP).[79] The Br-PCLA-PEG-

PCLA-Br was synthesized by conjugating 2-bromoisobu-

tyryl bromide with PCLA-PEG-PCLA. The pentablock



Figure 8. Schematic diagram of the sol–gel mechanism of the pH and temperature sensitiveblock copolymer solution. A) pH 7.4, 37 8C; B) pH 8.0, 37 8C; C) pH 7.4, 15 8C; D) pH 8.0, 15 8C.Reprinted with permission from ref.[75]. Copyright 2005 American Chemical Society.

Scheme 10. Structure of PAE-PCL-PEG-PCL-PAE.

572

copolymer was then prepared by polymerization of

sulfamethazine methacrylate monomers using Br-PCLA-

PEG-PCLA-Br as an ATRP macroinitiator. The molecular

weight distribution of the polymer was narrow relative to

that obtained from conventional radical polymerization.

Subsequently, OSM-PCGA-PEG-PCGA-OSM copolymers

were synthesized.[80] The OSM-PCGA-PEG-PCGA-OSM

hydrogel degraded at a faster rate than did the OSM-

PCLA-PEG-PCLA-OSM copolymer hydrogel.

Recently, copolymer hydrogels based on a basic poly(b-

amino ester) (PAE) were prepared.[81,82] PAE is known to be a

pH-sensitive, non-cytotoxic, biodegradable polymer, and

the positive charge of PAE facilitated an electrostatic

linkage with plasmid DNA (pDNA) at pH 7.2.[83] MPEG-PCL-

PAE block copolymers were synthesized by the Michael

addition polymerization of piperazine, hexan-1,6-diol

diacrylate (HDA), and MPEG acrylate.[81] Aqueous solutions

of the resulting copolymers exhibited a gel-to-sol transition

at pH values above 6.0, when the temperature was



increased. The phase diagram

could be tailored by varying the

MPEG molecular weight and PCL

block length. Gelation was related

to packing of the micelles upon

heating. Subsequently, a PAE-PCL-

PEG-PCL-PAE pentablock copoly-

mer was prepared by Michael

addition polymerization of 4,4-

trimethylene dipiperidine (TMDP),

PCL-PEG-PCL diacrylate, and

butane-1,4-diol diacrylate (BDA)

(Scheme 10).[82] The parent PCL-

PEG-PCL aqueous solution (20 wt.-%)

showed a sol–gel transition as a

function of temperature but

not pH. In contrast, the PAE-PCL-

PEG-PCL-PAE aqueous solution

at pH values above 6.0 underwent

a sol-to-gel transition in response

to both temperature and pH

changes (Figure 9). When the pen-

tablock copolymer was mixed with

insulin, the sol-to-gel transition

temperature was lowered because

of an ionic complex formed

between the polymer and insulin.

The PAE-PCL-PEG-PCL-PAE copoly-

mer hydrogel degraded in two

steps: first, fast degradation of

PAE was noted, followed by slow

degradation of the PCL-PEG-PCL

triblock copolymer. The PAE in

the hydrogel degraded completely

within 12 d, whereas 18 d were

required for degradation when the hydrogel was mixed

with insulin (a complex gel) in a pH 7.4 PBS solution. The

sol–gel transition of the polymer/insulin solution shifted

relative to the transition of the polymer solution alone.[82]

Thus, it was important to control this shift near physio-

logical conditions for practical applications.[84] The gel

window could be tailored by varying the PEG molecular

weight, PAE block length, PCL/PEG ratio, and concentration.

In addition, the degradation of pentablock copolymers

could be controlled by substituting PCLA for the PCL

block.[85] The PAE of the PAE-PCLA-PEG-PCLA-PAE hydrogel

degraded within 10 days compared to the 12 days for the

DOI: 10.1002/mabi.200900402


Figure 9. Sol–gel phase diagram of triblock and pentablock copo-lymer solutions at 20wt.-%. Reproduced with permission fromref.[81]. Copyright 2008 Elsevier.

Scheme 12. Structure of PAA-PEG-PAA.

PAE-PCL-PEG-PCL-PAE hydrogel, because of faster degrada-

tion of the PCLA block compared to the PCL block.

Subsequently, a series of pH/temperature-sensitive

multiblock copolymers based on poly(amino urethane)

(PAU) were reported.[86] The multiblock copolymers were

synthesized by polyaddition of HO-PCL-PEG-PCL-OH, bis-

1,4-(hydroxyethyl)piperazine (HEP), and 1,6-diisocynato

hexamethylene (HDI) (PCL-PEG-PCL-PAU)n (Scheme 11).

At pH values below 7.0, the polymer solution (20 wt.-%)

existed as a sol state over a temperature range of 0–80 8C. In

contrast, at pH 7.0 and above, the polymer solution

exhibited a sol-to-gel-to-aggregation transition upon heat-

ing. The sol–gel phase diagram could be controlled by

varying the hydrophobic/hydrophilic balance and block

length. A hydrogel formed quickly when the polymer

solution (20 wt.-%) was subcutaneously injected into rats.

The (PCL-PEG-PCL-PAU)n copolymers containing the hydro-

phobic PCL blocks showed incomplete solubility in water,

even at a relatively low pH. A double hydrophilic polymer

that did not contain a hydrophobic block dissolved easily in

water at low pH. Therefore, a series of pH/temperature-

sensitive multiblock copolymers based on the double

hydrophilic polymer were synthesized by polyaddition of

HO-PEG-OH, HEP, and HDI (PEG-PAU)m.[87] A 20 wt.-%

aqueous solution of (PEG-PAU)m showed a sol-to-gel-to-

sol transition in the pH range of 6.8–7.4 with increasing

temperature. Because of complete dissolution in water at

Scheme 11. Structure of [PCL-PEG-PCL-PAU]n.



low pH, the double hydrophilic multiblock copolymers may

be easily mixed with drugs or proteins, suggesting that

these copolymers are promising candidates for biomaterial

applications.

Unfortunately, it is difficult to control the molecular

weight and composition of PAU-related copolymers,

because of the presence of bifunctional monomers and

the high reactivity of HDI, resulting in multiblock

copolymers. A poly(amidoamine)-PEG-poly(amidoamine)

(PAA-PEG-PAA) triblock copolymer was synthesized by

Michael addition polymerization of PEG, TMDP, and 1,10-

decylene diacrylamide (DDA) (Scheme 12).[88] The composi-

tion of this copolymer was easily controlled, resulting in a

triblock copolymer. The PAA block formed a hydrophilic

block at relatively low pH (such as pH 3.0), but became

hydrophobic at higher pH (such as pH 7.4). The decrease

in pKa of the PAA-PEG-PAA polymer with increasing

temperature indicated that the PAA blocks became more

hydrophobic at higher temperatures. PAA-PEG-PAA aqu-

eous solutions underwent a sol-to-gel-to-condensed gel

transition with a very high viscosity of 43.6 kPa � s at 37 8Cand pH 7.4. At the condensed gel temperature, water was

squeezed from the gel matrix. The gelation process was

attributed to the self-assembly, hydrogen bonding, and

hydrophobic interactions of PAA blocks. The sol–gel

window was tailored by varying the polymer composition

and concentration.[89] The PAA-PEG-PAA copolymer

showed bioadhesive properties, suggesting potential for

applications involving mucosal surfaces. The mucoadhe-

sive properties were attributed to hydrogen bond interac-

tions between amide groups and the mucin, and the

interactions of positively charged PAA blocks with the

negatively charged sialic acid of mucin. In contrast to PAA-

PEG-PAA copolymers, which displayed a sol-to-gel transi-

tion with increasing temperature, concentrated PAE-PEG-

PAE triblock copolymers (30 wt.-%) in water showed a gel-

to-sol transition at higher temperatures and at pH> 6.4.[90]

PAE-PEG-PAE was synthesized by Michael addition poly-

merization of PEG diacrylate, TMDP, and HDA. The resulting

polymer could be easily dissolved in water at a

relatively low pH, because of its

doubly hydrophilic character. The

sol–gel transition point was

adjusted by changing the composi-

tion. This polymer also

showed potential for bioadhesive

applications.



574

Biomedical Applications of InjectableBiodegradable Hydrogels

Biodegradable polymeric hydrogels that are responsive to

temperature or pH/temperature changes have been widely

Table 1. Various biodegradable injectable hydrogel systems and thei

Hydrogel types Polymer

Temperature-sensitive PEG-PLLA-PEG

PEG-PLGA-PEG

PLGA-PEG-PLGA

(ReGel1)

PLGA-g-PEG/PEG-g-

PLGA

MPEG-PCL

Cholesterol end-capped

star PEG-PLLA

PEG/PPG/PHB

Polyphosphazenes

C/GP

pH/temperature-sensitive OSM-PCLA-PEG-PCLA-OSM

PAE-PCL-PEG-PCL-PAE

(PCL-PEG-PCL-PAU)n

(PEG-PAU)m

PAA-PEG-PAA

PAE-PEG-PAE



used for biomedical applications, such as drug/protein

delivery and tissue engineering. Aqueous polymer solu-

tions may be loaded with drugs, proteins, or cells at a

specific temperature before injection into particular sites in

the body. Once formed, the hydrogels act as drug delivery

r applications.

Delivery/applications References

FITC-labeled dextran [9]

Spironolactone, ketoprofen [91]

Plasmid DNA [92]

Plasmid TGF-b1 [93]

FITC [94]

Insulin [95]

GLP-1 [96]

Testosterone [97]

Indomethacin, 5-fluorouracil [98]

Ganciclovir [99]

Levonorgestrel, interleukin-2,

ceftazidime

[100–102]

Paclitaxel [23]

Insulin [103]

Chondrocyte

rBMSC, dexamethasone [104]

FITC-BSA, BSA [105,106]

L929 cells [46]

BSA [48]

FITC-dextran, HSA [107]

Doxorubicin, paclitaxel [108,112]

FITC-albumin [109]

Pancreatic islets, hepatocytes [110,111]

Camptothecin, insulin [113,114]

Bovine chondrocytes [115]

Paclitaxel [116]

hMSCs [117]

Insulin [82,118]

Paclitaxel [86]

Chlorambucil [87]

Flurbiprofen [90]

Lidocaine [89]

DOI: 10.1002/mabi.200900402


matrices or cell growth depots. Table 1 summarizes the

biomedical applications of the biodegradable injectable

hydrogels.

Temperature-Sensitive Block Copolymer Hydrogels

Thermogelling PEG-PLLA-PEG hydrogels were investigated

for the sustained release of fluorescein isothiocyanate

(FITC)-labeled dextran, which was incorporated into poly-

mer solutions at 45 8C, and the drug-loaded polymer

solutions were changed to gels by cooling to body

temperature. The release rate of dextran was influenced

by polymer concentration. With a 35 wt.-% polymer gel,

dextran was released at a constant rate over the course of

12 days, without a burst release, in contrast with the burst

effect observed for 23 wt.-% polymer gels.[9]

To study the influence of hydrophobicity and hydro-

philicity on sustained release using a PEG-PLGA-PEG

triblock copolymer hydrogel, spironolactone and ketopro-

fen were employed as drug models.[91] The hydrophilic

ketoprofen was released over two weeks with a first-order

release profile, whereas hydrophobic spironolactone

required two months for release and yielded an S-shaped

release trace. A diffusion mechanism was proposed for

release of the hydrophilic drug, whereas diffusion at the

first stage, followed by degradation at later stages, was

the mechanism envisaged to describe the release of the

hydrophobic drug. pDNA was released from a PEG-PLGA-

PEG matrix at an approximately constant rate (zero order)

over two weeks, and the drug half life was five days.[92]

When a hydrogel mixed with luciferase pDNA was applied

to skin wounds of CD-1 mice, the expression of luciferase

reached a maximum at 24 h and then dropped to 94% at

72 h. The supercoiled structure of the released pDNA was

preserved, although a small quantity of linear pDNA was

observed. A plasmid TGF-b1-loaded PEG-PLGA-PEG hydro-

gel was used to accelerate wound healing in diabetic

mice.[93] Reepithelialization was complete at day 9 using

hydrogels containing TGF-b1, at day 11 employing a

commercial wound dressing (Humatrix) that contained

TGF-b1, and at day 14 in the absence of treatment.

Fibroblast proliferation and collagen organization were

significantly enhanced by hydrogel treatment relative to

treatment with TGF-b1-loaded Humatrix. These results

suggest that PEG-PLGA-PEG hydrogels could be used as a

gene delivery system for treatment of skin disorders and

wound healing. A hydrogel mixed with FITC was found to

prolong drug exposure in the bladder of rats.[94] The FITC-

loaded hydrogel showed sustained release for up to 24 h

after instillation, whereas all administered free FITC was

observed in the urine of rats after 8 h.

The in vitro release of human insulin from PLGA-PEG-

PLGA (ReGel) systems proceeded at a constant rate over two

weeks, without an initial burst effect.[95] The presence of



0.2% (w/v) zinc could enhance the insulin release rate, and

almost 90% of the insulin was released over the course of

two weeks. An in vivo study revealed that a steady amount

of insulin was secreted from the ReGel/0.2 wt.-% Zn-insulin

depot within a period of two weeks. A related study showed

that the incretin hormone glucagon-like peptide-1 (GLP-1)

was released from ReGel in vitro and in vivo.[96] The in vitro

study of GLP-1/ReGel showed a release profile over five

days, whereas the release profile of zinc-complexed GLP-1/

ReGel (ZnGLP-1/ReGel) exhibited a zero-order release profile

over two weeks, without an initial burst. After a single

injection of ZnGLP-1/ReGel into Zucker diabetic fatty (ZDF)

rats, plasma GLP-1 was maintained at levels that were

significantly higher than the control group over the course

of two weeks. In addition, plasma insulin levels increased

and the blood glucose levels fell.

Testosterone was released over the course of three

months from PLGA-PEG-PLGA systems.[97] The drug release

rates were affected by drug concentration, solvent compo-

sition, and composition of the copolymer. The release of

indomethacin and 5-fluorouracil from PLGA-PEG-PLGA

systems was also reported.[98] Hydrophilic 5-fluorouracil

and hydrophobic indomethacin were secreted from the

hydrogel over five days and one month, respectively, and

the release rate depended on copolymer composition.

Ganciclovir-loaded PLGA microspheres were dispersed in a

PLGA-PEG-PLGA gel.[99] The microspheres/hydrogel showed

a slower release rate compared to that from PLGA micro-

spheres alone. The release of levonorgestrel, interleukin-2,

and ceftazidime from a PLGA-PEG-PLGA hydrogel was

reported.[100–102] ReGel was used to release a hydrophobic

drug, paclitaxel (ReGel/paclitaxel: OncoGel),[23] over the

course of 50 d. The in vitro release study showed diffusion-

controlled release within the first two weeks, followed by

combined diffusion/polymer degradation.

A PLGA-g-PEG/PEG-g-PLGA hydrogel was used for the

sustained release of insulin.[103] After a single injection of

insulin-loaded hydrogel, blood glucose levels could be

adjusted from 5 to 16 days in diabetic rats, depending on

polymer composition. Also, a chondrocyte-loaded PLGA-g-

PEG gel was used to repair an articular cartilage defect.

Poly(N-isopropyl acrylamide)-co-acrylic acid/hydroxyapa-

tite collagen sponges containing chondrocytes were used as

a control. The cartilage defect was completely repaired

using the PLGA-g-PEG hydrogel, in contrast with persis-

tence of control defects. The superior efficacy of cartilage

defect repair was attributed to the biodegradability of the

PLGA-g-PEG.

The in vivo osteogenic differentiation of rat bone marrow

stromal cells (rBMSC) was studied.[104] Histological analysis

demonstrated that in situ formed MPEG-PCL hydrogels

containing rBMSC and dexamethasone were biocompatible

and enhanced bone formation. The in vitro release of FITC-

labeled bovine serum albumin (FITC-BSA) from the MPEG-



576

PCL hydrogel showed a sustained release profile for more

than 20 days.[105] Although in vivo release of FITC-BSA was

sustained for 30 days, an initial burst was observed. In

addition, MPEG-PCL wafers could be prepared as an

implantable material using a direct compression

method.[106] BSA was released from the wafers over the

course of 30 days with an initial burst release.

L929 cells encapsulated in cholesterol end-capped star

PEG-PLLA hydrogels (10 and 20 wt.-%) were viable and

proliferated in three dimensions within the hydrogels,[46]

indicating that the polymer could be used as an injectable

cellular scaffold. The PEG/PPG/PHB hydrogel is a promising

candidate for the controlled release of protein.[48] The in

vitro release of BSA was controllable over 70 d. Diffusion

control was proposed as the release mechanism during the

first stage of release, and the mechanism at the later stages

was proposed to be erosion control.

Polyphosphazenes were used for the sustained release of

FITC-dextran and human serum albumin (HSA) over the

course of two weeks.[107] The release of FITC-dextran was

dependent on polymer concentration. The solubility of

doxorubicin (DOX) was enhanced in polyphosphazene

hydrogels.[108] DOX was released at a controlled rate from

the polyphosphazene hydrogel over 20 d and the release

rate was affected by gel strength. The antitumor activity of

DOX in the mouse lymphoblast cell line P388D1 was

maintained over the course of one month. The release of

FITC-albumin from polyphosphazene hydrogels was con-

trolled using chitosan,[109] and was sustained over two

months without an observed initial burst in the presence of

chitosan, in contrast with the observed release over one

month without chitosan. The extension of release time was

attributable to formation of an ionic complex between

chitosan and FITC-albumin. The polyphosphazene hydrogel

has also been used to entrap pancreatic islets.[110] In

comparison with both rat islets entrapped in other

hydrogels, and free islets, rat islets in the polyphosphazene

hydrogel retained higher cell viability and insulin produc-

tion over a 28-day culture period. In subsequent work,

polyphosphazene hydrogels were used to encapsulate

hepatocytes as spheroids or single cells.[111] Over a

28-day culture period, the spheroid hepatocytes main-

tained a higher viability and produced albumin and urea,

whereas single hepatocytes reduced the level of albumin

secretion from the hydrogel. These results suggest that the

hydrogels could be used as a three-dimensional cell culture

system. More recently, paclitaxel within covalently con-

jugated polyphosphazene gels showed sustained release

over one month at pH 7.4, and three days at pH 6.8.[112] The

in vitro evaluation of antitumor activities against several

cancer cell lines indicated that paclitaxel was released from

the hydrogel without inhibiting tumor growth. However,

the paclitaxel-conjugated polyphosphazene hydrogel

showed a higher level of antitumor activity over one



month compared with the control, after a single intratu-

moral injection of the hydrogel into HSC-45M2 human

gastric cancer cell-containing nude mice.

C/GP was used as a hydrogel for the sustained release of

camptothecin.[113] The release profile showed zero-order

release kinetics within the first four weeks. In contrast with

the blank C/CP, tumor growth was significantly delayed

after injection of C/GP-containing camptothecin. Insulin

was mixed with G/GP and released over 350 h. At a fixed

concentration of C (2.5 v/w), the higher GP content resulted

in a faster release of insulin, due to the higher mobility of

the protein in the gel. The diffusion controlled process was

attributed to the release mechanism.[114] A chitosan-g-

pluronic (CP) system was used for injectable cell delivery to

generate cartilage.[115] When bovine chondrocytes were

encapsulated in a CP hydrogel, cell viability and synthesized

glycosaminoglycan content increased after 28 days of cell

culture.

pH/Temperature-Sensitive Block CopolymerHydrogels

OSM-PCLA-PEG-PCLA-OSM hydrogels with good biocom-

patibility were used for the controlled release of pacli-

taxel.[116] An in vitro release (pH 7.4, 37 8C) study showed

sustained release at a constant rate over 20 d. An in vivo

study was carried out on C57BL/6 male mice. Good anti-

tumor activity was observed over the course of two weeks

after subcutaneous injection of the paclitaxel-loaded

copolymer hydrogel into tumor-bearing mice. After two

weeks of treatment, the tumor volume of saline-treated

mice was about 17 cm3, whereas that of paclitaxel/

hydrogel-treated mice smaller than 7 cm3. The anti-tumor

effect depended on the concentration of paclitaxel. At the

paclitaxel concentration of 1 mg �mL�1, the body weight

was constantly maintained for two weeks while the body

weight decreased during the first six days and slightly

increased. The release of paclitaxel from OSM-based

hydrogels was controlled using PLGA-PEG-PLGA instead

of PCLA-PEG-PCLA.[77] In comparison with the release profile

of paclitaxel from the OSM-PCLA-PEG-PCLA-OSM hydrogel,

the release of paclitaxel from the OSM-PCGA-PEG-PCGA-

OSM hydrogel was faster, because of more rapid degrada-

tion of the PCGA block. A OSM-PCLA-PEG-PCLA-OSM

solution containing human mesenchymal stem cells

(hMSCs) and recombinant human bone morphogenetic

protein-2 (rhBMP-2) was injected into the backs of mice.[117]

After seven weeks, mineralized tissue with high levels of

alkaline phosphate activity had formed, indicating that this

material could be used as an injectable scaffold for bone

tissue engineering.

pH/temperature-sensitive PAE-PCL-PEG-PCL-PAE hydro-

gels were employed for the controlled release of insulin.[82]

The polymer exhibited as a solution at pH< 6, thus, it was

DOI: 10.1002/mabi.200900402


Figure 10. Insulin release experiment in vivo. In the insulin-onlygroup, a 200mL insulin solution of 0.25mg �mL�1 (in PBS buffer(pH 7.4) is administered by subcutaneous injection (0.05mg ofinsulin for each rat)). In the insulin PCL-PEG-PCL gel group, 200mLof a solution (5mg �mL�1 of insulin in PCL-PEG-PCL solutions of25wt.-%) at pH 7.0 and 10 8C is subcutaneously injected (1mg ofinsulin for each rat). In the complex gel group, 200mL of thecomplexed insulin solution (5mg �mL�1 in PAE-PCL-PEG-PCL-PAEsolutions of 25wt.-%) at pH 7.0 and 10 8C is subcutaneouslyinjected (1mg of insulin for each rat). (Male SD rats, error barsrepresent the standard deviation (n¼ 5). Reproduced with per-mission from ref.[81]. Copyright 2008 Elsevier.

easy to mix the polymer and insulin at pH 3.0–4.0 at 2 8C. An

in vitro study carried out at pH 7.4 and 37 8C showed a

constant release rate profile for up to 20 days. After a single

injection of the PAE-PCL-PEG-PCL-PAE/insulin solution into

SD rats, the serum insulin concentration was sustained for

one month and plasma insulin levels were maintained at a

constant level over 15 days, without an initial burst

(Figure 10). In contrast, a 1-day duration of plasma insulin

level and a marked initial burst were observed when

the PCL-PEG-PCL/insulin mixture was employed. The

physiological effects of pentablock copolymer/insulin

injection were investigated in diabetic rats.[118] After a

single injection into a diabetic rat, steady blood glucose

levels were maintained for more than one week, without a

decrease in body weight. The blood glucose levels and body

weight showed a dose dependence of insulin loaded into

the polymer. The sustained release of insulin was related to

the ionic interactions between partial positive charges in

PAE blocks and negative charges in insulin at

physiological pH. The sustained release of insulin was

attributed to degradation of the PAE blocks, suggesting that

a degradation-controlled process was the major mechan-

ism of insulin release.

The (PCL-PEG-PCL-PAU)n multiblock hydrogel was used

for sustained release of paclitaxel over the course of one



month, and chlorambucil was released from the (PEG-

PAU)m multiblock hydrogel for 10 days under physiological

conditions. For these polymers, drug encapsulation was

easily performed at pH 5.0–6.0 and 2 8C because of their

solubility character.[86,87] PAE-PEG-PAE and PAA-PEG-PAA

hydrogels may have potential for drug delivery at mucosal

surfaces.[89,90] Lidocaine and flurbiprofen were released

from PAE-PEG-PAE and PAA-PEG-PAA hydrogels over the

course of one day, indicating that these hydrogels could be

used as drug carriers.

Conclusions and Perspectives

Significant progress has been achieved in the development

of injectable biodegradable polymeric hydrogels, and each

hydrogel system has special intrinsic properties (including

gel strength, pH after degradation, and degradation rate)

that may be appropriate for a particular application. In this

review, the characteristics, sol–gel mechanisms, and

biomedical applications (such as drug/cell delivery and

tissue engineering) of temperature- and pH/temperature-

sensitive hydrogels are summarized. Some challenges

remain for improving the applicability of hydrogels.

The first challenge in the development of practical

injectable applications of thermosensitive block copolymer

hydrogels is administration. The risk of the syringe clogging

upon injection can be addressed by modulating the gelation

temperature and lowering the polymer concentration.

Slowly degrading polymers can be used for long-term

applications in drug/cell delivery and tissue engineering,

thus, the lifetime of the gel prior to degradation must

be considered. For controlled drug release, an initial burst

release is a limiting factor, especially for low-molecular-

weight drugs, hydrophilic drugs, and proteins. The require-

ment of storage in solution for commercial use may cause

degradation, thus, the reconstitution problem should

be considered. The products generated from the degrada-

tion of some polymers, such as PLLA and PLGA, may lower

the pH of the surrounding environment and should,

therefore, be considered. Such degradation products may

cause inflammation at the injection site or damage

incorporated proteins/cells.

pH/temperature-sensitive block copolymer hydrogels

show several advantages over thermosensitive block

copolymer hydrogels, such as the absence of clogging

during injection, which allows facile injection into deep

sites in the body, avoidance of the local low pH environ-

ment caused by degradation, which protects proteins/cells

from damage, and the ease of handling and storage. In

particular, cationic block copolymer hydrogels can form

ionic linkages with the anionic proteins or pDNA at

physiological conditions, which result in sustained release

without an initial burst. The sustained release of the low-


KHANHNGUYEN

Highlight


578

molecular-weight and hydrophilic drugs may be obtained

by ionically bonding with the polymers. The pentablock

copolymers possess complicated structures that may make

FDA approval difficult, thus, well-defined pH/temperature-

sensitive copolymers should be considered. In addition,

the degradation rate of such well-defined copolymers

should be optimal for the sustained release of proteins over

one month.

For both kinds of the above mentioned block copolymer

hydrogels, appropriate time for clearance of the hydrogels

from the body is quite important. The degradation depends

on polymer composition, crystallinity of the polymer, and

topological structure. This understanding may help to

design hydrogels with fine tuning of the degradation rate,

which results in the corresponding release rate. The

changes of the gel phase and gel strength after mixing

with drugs, cells, or proteins should be a challenge. The

integration of drugs may be considered for potential

applications. Biodistribution, elimination routes and the

effect of degradation products on organs, are also of

concern. For tissue engineering and cell growth, cell

adhesion to the gels, molecular-level characteristics of

the tissue/polymer interface, effects of the degradation

products on cells, and the effect of hydrogels on histogen-

esis, should be further investigated.

Acknowledgements: The authors thank Dr GuangJin Im for hiscomments on this paper. This research was financially supportedby the MEST and NRF (20090093631)

Received: November 2, 2009; Revised: December 15, 2009;Published online: March 1, 2010; DOI: 10.1002/mabi.200900402

Keywords: biodegradable; block copolymers; hydrogels; inject-able; pH/temperature-sensitive; temperature-sensitive

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