Injectable and tunable poly(ethylene glycol) analogue hydrogels based on poly(oligoethylene glycol methacrylate)

Article abstract of DOI:10.1039/c3cc48514e

Injectable PEG-analogue hydrogels based on poly(oligoethylene glycol methacrylate) have been developed based on complementary hydrazide and aldehyde reactive linear polymer precursors. These hydrogels display the desired biological properties of PEG, form

Full text of DOI:10.1039/c3cc48514e

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Injectable and tunable poly(ethylene glycol)
analogue hydrogels based on poly(oligoethylene
Cite this:DOI: 10.1039/c3cc48514e
glycol methacrylate)†
Received 7th November 2013,
Accepted 24th January 2014
Niels M. B. Smeets,‡ Emilia Bakaic,‡ Mathew Patenaude and Todd Hoare*
DOI: 10.1039/c3cc48514e
www.rsc.org/chemcomm
Injectable PEG-analogue hydrogels based on poly(oligoethylene glycol potential applications. POEGMA is synthesized via simple free
methacrylate) have been developed based on complementary hydra- radical copolymerization6 and has been demonstrated to serve
zide and aldehyde reactive linear polymer precursors. These hydrogels as an effective PEG analogue, exhibiting corresponding non-
display the desired biological properties of PEG, form covalent net- immunogenic, non-cytotoxic and protein repellent properties to
works in situ following injection, and are easily modulated for improved PEG.7 A number of POEGMA-based hydrogels have been reported
control over their functionality and physiochemical properties.
to-date;8 however, none are covalently cross-linked while also being
injectable and degradable in vivo, significantly limiting their
Poly(ethylene glycol) (PEG) hydrogels have been extensively studied potential clinical application.
as matrices for the controlled release of therapeutics and as scaffolds
Our PEG-analogue hydrogels are based on hydrazide and alde-
for promoting tissue regeneration.1 PEG hydrogels are hydrophilic, hyde functionalized POEGMA precursors that rapidly form a hydro-
non-cytotoxic and non-immunogenic and can effectively mask the gel network via reversible hydrazone bond formation upon
material from the host’s immune system.2 A significant drawback of co-extrusion9 (see Scheme 1). The injectable, in situ gelling nature
PEG, however, is that the polymer lacks chemical versatility given of this system is useful to circumvent many of the issues concerning
that functionalization is limited to the hydroxyl chain end(s).3 PEG surgical implantation of bulk hydrogel-based materials. The
hydrogels are predominantly synthesized via step-growth polymer- hydrazide-functionalized POEGMA precursors (PO100Hy) were syn-
ization of complimentary a,o-functionalized PEG precursors using a thesized by copolymerizing OEGMA475 (EO repeat units, n = 8–9)
range of different chemistries.4 Although step-growth polymer- with acrylic acid (AA) and subsequent post-polymerization modifica-
ization minimizes network non-idealities, chemical modification tion using EDC chemistry with a large excess of adipic acid
of linear PEG precursors is limited by the number of available dihydrazide. The aldehyde-functionalized POEGMA precursors
reactive end groups. As such, there is increasing interest in (PO100Ay) were synthesized by copolymerizing OEGMA475 with a
polymers with similar (biological and physicochemical) proper- functional acetal monomer (N-(2,2-dimethoxyethyl)methacrylamide,
ties that can be easily synthesized with improved control over the DMEMAm) and subsequently converting the acetal to the corre-
polymer functionality.5
sponding aldehyde by acid-catalyzed hydrolysis (Scheme 1).
Hereto, we prepared PEG-analogue hydrogels based on
poly(oligoethylene glycol methacrylate) (POEGMA) that exhibit all
of the desirable protein and cell-repellent properties of conven-
tional PEG hydrogels while also being injectable, degradable, and
easily chemically and mechanically tunable. The approach enables
facile preparation of hydrogels with potential utility as injectable
tissue engineering matrices, local drug delivery vehicles for small
molecules or macromolecules, or joint lubricants, among other
Department of Chemical Engineering, McMaster University, 1280 Main Street West,
Hamilton, Ontario, Canada L8S 4L7. E-mail: hoaretr@mcmaster.ca
† Electronic supplementary information (ESI) available: Detailed experimental
procedures, chemical characterization and in vitro toxicity of the POEGMA
precursors, as well as hydrogel swelling kinetics, degradation, rheology and
histology. See DOI: 10.1039/c3cc48514e
Scheme 1 Preparation of injectable PEG analogue hydrogels.
‡ These authors contributed equally to this work.
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The degree of functionality of the PO100Hy and PO100Ay precursors PO100A30 confirmed that the molecular weight distribution (MWD)
was controlled by varying the AA and DMEMAm content from y = 20 of the degradation products is directly analogous to the combined
to 40 mol% (ESI,† Table S1 and Fig. S1). The number-average MWDs of both precursor polymers (see ESI,† Fig. S3). Consequently,
molecular weight of the precursor polymers was controlled to be the PO100Hy and PO100Ay polymers represent both the hydrogel
lower than 20 ꢀ 103 g molꢁ1, well below the renal clearance limit of precursors as well as the hydrogel degradation products. MTT assay
40–50 ꢀ 103 g molꢁ1 to facilitate polymer elimination following gel results following exposure of 3T3 mouse fibroblasts to PO100Hy and
degradation.10
PO100Ay confirm that these materials do not impart any significant
POEGMA hydrogels were prepared by co-extruding PO100Hy and in vitro toxicity up to concentrations of 2000 mg mLꢁ1 (ESI,† Fig. S4),
PO100Ay solutions in 10 mM PBS using a double-barrel syringe. suggesting that both the precursors and degradation products of the
Gelation occurs over time frames ranging from several hours (B8 h) hydrogels are compatible with cells.
to a few minutes (o10 min) depending on functional group content
The degradation rate of the POEGMA hydrogels is governed by
and polymer concentration (ESI,† Table S2). The hydrogels swell in the cross-link density and, by extension, the degree of function-
PBS following preparation (Fig. 1E–J), indicating high water affinity. alization and concentration of the PO100Hy and PO100Ay precursors
For POEGMA hydrogels prepared with precursors containing 25, 30 (Fig. 1B). Hydrogels prepared with precursors at low concentra-
or 40 mol% functional groups (both hydrazide and aldehyde), the tions (o125 mg mLꢁ1) or with low degrees of functionalization
equilibrium mass-based swelling ratio (Qm) is reached within 30 h (o30 mol%) degrade in o1 min in 100 mM HCl, while hydrogels
(ESI,† Fig. S2) and decreases systematically with both the degree of prepared with precursors at a high concentration (200 mg mLꢁ1
)
chain functionalization and the concentration of precursor poly- or with a high degree of functionalization (40 mol%) degrade
mers used to prepare the hydrogel (Fig. 1A). This result suggests that significantly slower, requiring approximately 5 hours to fully
facile tuning of the hydrogel water content at equilibrium is possible degrade under acid-catalyzed conditions (Fig. 1B). Long-term
based on the types and concentrations of precursor polymers used incubation showed that the hydrazone cross-linked POEGMA
(see ESI†).
hydrogels (150 mg mLꢁ1 and 30 mol% reactive hydrazide and
The injectable POEGMA hydrogels are cross-linked through the aldehyde groups) are stable for at least 5 months under physio-
formation of dynamic hydrazone bonds, which are reversible in an logical conditions in vitro but degrade within 4 weeks in vivo
aqueous environment. Aqueous size exclusion chromatography of following subcutaneous injection in BALB/c mice.
the degradation products of a hydrogel prepared from PO100H30 and
The PO100Hy and PO100Ay precursors yield hydrogels with
elastic moduli that can be tuned depending on the number of
reactive functional groups as well as the precursor concen-
tration. G0 values range from 0.23 ꢂ 0.02 kPa (150 mg mLꢁ1
,
25 mol%) to 8.0 ꢂ 1.0 kPa (150 mg mLꢁ1, 40 mol%) (Fig. 1C and
D and ESI,† Fig. S5 and S6). For comparison, the G0 of conven-
tional PEG hydrogels prepared from multi-arm and di-functional
PEG precursors at comparable concentrations typically ranges
from 0.25 to 6.0 kPa,4b,c, f analogous to the range reported here.
However, the modulus of these POEGMA hydrogels can be tuned
independently of polymer concentration (and thus hydrogel
osmotic pressure) if desired by modifying the precursor func-
tionality, not possible using conventional PEG chemistry.
The biointerfacial properties of the injectable POEGMA hydro-
gels (prepared at 150 mg mLꢁ1 with 30 mol% functional groups)
were evaluated using protein adsorption (Fig. 2A and B) and cell
adhesion (Fig. 2C–E) assays. POEGMA hydrogels were incubated
with bovine serum albumin (BSA) and fibrinogen (Fib) at varying
concentrations. BSA and Fib adsorption to the POEGMA hydrogel is
maintained below 100 and 500 ng cmꢁ2 respectively even upon
exposure to a 2000 mg mLꢁ1 protein solution (Fig. 2A and B),
comparable to adsorption values reported for conventional PEGy-
lated surfaces (see ESI,† Table S4).11 Note that protein absorption is
also likely to occur for POEGMA hydrogels, suggesting that the true
protein adsorption on the hydrogel surface is likely even lower than
reported in Fig. 2A and B. Furthermore, negligible adhesion of
fibroblast cells (noted to adhere particularly strongly to many
biomaterials due to their inherent ability to produce extracellular
matrix)12 is observed to the POEGMA hydrogels, with 0 ꢂ 1 cells per
mm2 (n = 6) observed on POEGMA hydrogel interfaces (Fig. 2D and
ESI,† Fig. S7) compared to 2664 ꢂ 100 cells per mm2 (n = 6) for
Fig. 1 Physiochemical characterization of injectable POEGMA hydrogels.
(A) Bottom point = initial swelling and top point = equilibrium swelling
(n) 25 mol%, (K) 30 mol% and (,) 40 mol% hydrazides. (B) Degradation
kinetics at 30 mol% functionality in 100 mM HCl (J) 100 mg mLꢁ1, (K)
125 mg mLꢁ1, (K) 150 mg mLꢁ1, (K) 175 mg mLꢁ1, (K) 200 mg mLꢁ1 and (n)
150 mg mLꢁ1 and 40 mol%. (C and D) Elastic storage modulus as a function
of the precursor functionalization (C) and concentration (D). (E–J) Photo-
graphs of the unswollen and swollen hydrogels (grid = 0.5 mm ꢀ 0.5 mm).
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in the hydrogel (Fig. 1B); in this sense, hydrogel clearance kinetics
in vivo may also be engineered.
In conclusion, we have prepared injectable, hydrazone cross-
linked hydrogels based on POEGMA that exhibit the same
favourable biological properties of conventional PEG-based
hydrogels (i.e. facilitating minimal protein adsorption, no sig-
nificant cellular adhesion, and no significant chronic inflam-
matory responses in vivo) while offering the significant
advantages of facile synthesis, rapid in situ gelation following
injection, tunable mechanics, tunable degradation times, and
excellent control over chemical composition and functionaliz-
ability. These results suggest the potential of these POEGMA-
based hydrogels as a platform for the design of engineered
hydrogels for a variety of biomedical applications now served by
conventional PEG hydrogels.
The NSERC 20/20 Ophthalmic Research Network is grate-
fully acknowledged for financial support.
Fig. 2 Biological properties of injectable POEGMA hydrogels (based on
0
PO100H30 and PO100A30 15 wt%). (A and B) Adsorption of BSA (A) and
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