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Research Paper
Influence of physical and mechanical properties
of amphiphilic biosynthetic hydrogels on long-term
cell viability
Finosh Gnanaprakasam Thankam, Jayabalan Muthu
n
Sree Chitra Tirunal Institute for Medical Sciences and Technology, Polymer Science Division, BMT Wing,
Thiruvananthapuram 695012, Kerala State, India
article info
Article history:
Received 5 March 2014
Accepted 9 March 2014
Available online 12 April 2014
Keywords:
Biosynthetic hydrogels
Mechanical properties
Long term viability
Amphiphilicity
Cross linking density
abstract
Maintaining the mechanical properties of biofunctional hydrogels of natural resources for
tissue engineering and biomedical applications for an intended period of duration is a
challenge. Though anionic polysaccharide alginate has been hailed for its excellent
biomimetic characters for tissue engineering, it usually fails in load bearing and other
dynamic mechanical environment. In this paper this issue was addressed by copolymeriz-
ing alginate with the biocompatible and mechanically robust synthetic biodegradable
polyester and crosslinking with polyethylene glycol diacrylate (PEGDA) and vinyl co-
monomers, 2-hydroxy ethyl methacrylate (HEMA), methyl methacrylate (MMA) and N N'
methylene bis acrylamide (NMBA) to form three hydrogels. All three hydrogels were
amphiphilic, hemocompatible and non-cytotoxic. These hydrogels exhibited appreciable
water holding capacity. Comparatively, hydrogel prepared with PEGDA–NMBA crosslinkers
displayed larger pore size, increased crosslinking, higher tensile strength and controlled
degradation. With appreciable swelling and EWC, this hydrogel elicited better biological
responses with long-term cell viability for cardiac tissue engineering.
&2014 Elsevier Ltd. All rights reserved.
1. Introduction
Even though the natural hydrogel materials exhibit better
biocompatibility and cell affinity for tissue engineering appli-
cations than synthetic polymers, they are less versatile when
considering properties like mechanical strength and variable
degradation rate. Copolymerization of a mechanically robust
synthetic polymer with a biocompatible natural polymer can
overcome most of these difficulties. The scaffolds made from
such a copolymer can maintain the cell phenotype and
modulate its function because the synthetic materials can
greatly enhance properties of natural polymer (Li and Guan,
2011). An ideal scaffold for tissue engineering should possess
the structural characteristics of synthetic materials and the
biofunctional characteristics of natural materials.
The advent of biosynthetic hybrid hydrogel is highly
relevant for tissue engineering scaffolds, especially to that
of heart, as it can solve several mechanical and biological
http://dx.doi.org/10.1016/j.jmbbm.2014.03.010
1751-6161/&2014 Elsevier Ltd. All rights reserved.
n
Corresponding author. Tel.: þ0091 471 2520212; fax: þ0091 471 234814.
E-mail address: mjayabalan52@gmail.com (J. Muthu).
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122
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issues arising during the individual use of synthetic or
natural polymers. Scaffold materials of natural origin like
alginate were reported to be effective for mediating cell-
signaling mechanisms for cardiac repair and regeneration.
Despite their poor mechanical properties, the inflammatory
potential is high which depends on the type, chemistry and
source. On the other hand, synthetic cardiac scaffold materi-
als can be customized to get a wide range of tunable proper-
ties in terms of degradation, mechanical strength and
durability. Still their inherent cellular interactions are limited;
however this can be improved by functionalization with
biological moieties Novakovic et al., 2010).
Among the natural biomaterials, the seaweed-derived
polysaccharide alginate attracted the attention of many
tissue engineers mainly due to its inert chemistry that allows
interactions with mammalian cells. Despite their mechanical
issues, many researchers explored alginate for tissue engi-
neering especially for cardiac applications (Andersen et al.,
2012;Rosellini et al., 2009). Alginate is composed of 1, 4-linked
β-D-mannuronate and 1, 4-linked α-L-guluronate units that
forms a linear polymer which can undergo gelling by chelat-
ing calcium ions. The calcium alginate gel is immunologically
inert and non-digestible by mammalian cells. Calcium algi-
nate gels can degrade in the biological systems by the gradual
diffusion of calcium ions and exchange with monovalent ions
from the medium. The degradation products are non-toxic
and are excreted through the urine (Novikova et al., 2006).
Moreover, alginate in its hydrogel form closely imitates
native glycocomponents of ECM (Leor and Cohen, 2004).
The unsaturated linear polyester poly(propylene fumarate)
(PPF) can be modified or crosslinked through its double bonds.
Therefore, it is an ideal choice for various biomedical applica-
tions. PPF degrades by simple hydrolysis of the ester bonds;
the degradation products primarily fumaric acid and propy-
lene glycol are proven to be non-toxic and can enter the
mitochondrial TCA cycle. The unsaturation present in the PPF
allows the crosslinking of the polymer into a covalent poly-
mer network. These networks can be designed with a wide
range of controlled properties as needed for cardiac tissue
engineering (Kasper et al., 2009;Finosh and Jayabalan, 2012).
Linking the biological polymer with structurally versatile
synthetic polymers will pave a way for the construction of a
crosslinked network with a balanced physiochemical and
biomechanical properties and a controlled degradation pro-
file, which is a requisite for the in vitro engineering of organ
parts. Hydrogels are composed mainly of hydrophilic polymer
network that offers a strong rapport for water while their
physical or chemical crosslinking ensures the association of
water by preventing the dissolution. Nevertheless, the water
penetrates into the polymer network and allows swelling.
The high water content offers them excellent biocompatibil-
ity by which hydrogels have become as an ideal choice for
tissue engineering, drug delivery, and bio-nanotechnology
applications (Peppas et al., 2006).
Based on the background we aimed at the synthesis of
copolymer (AP) hydrogel scaffolds using alginate and PPF. In
order to enhance the mechanical properties and amphiphilic
character for cell growth, the AP copolymer was modified
with polyethylene glycol diacrylate (PEGDA) and vinyl mono-
mers to form hydrogels. The present paper deals with studies
on the influence of physical and mechanical properties
of amphiphilic biosynthetic hydrogels on long-term cell
viability.
2. Methodology
2.1. Materials
Sodium alginate [guluronic acid (39%) and mannuronic acid
(61%) from brown algae, medium viscosity, Product no. A2033],
sodium chloride, maleic anhydride, calcium chloride, L-ascorbic
acid, ammonium per sulfate, polyethylene glycol diacrylate
(PEGDA), 2-hydroxy ethyl methacrylate (HEMA), methyl metha-
crylate (MMA) and N N' methylene bis acrylamide (NMBA) were
obtained from Sigma-Aldrich, Spruce Street, St. Louis, USA.
Sodium acetate, sodium hydroxide etc. were supplied by Merck
specialties Pvt. Ltd, Mumbai, India. 1-2 propylene glycol and
morpholine were provided by SD fine chemicals India Ltd. The
mouse fibroblast cells (L929) were purchased from National
Centre for Cell Science, Pune, India for cell culture studies.
2.2. Preparation of chemically crsslinked and mechanically
favorable biosynthetic Hydrogel scaffolds
Hydroxyl terminated-poly (propylene fumarate) (HT-PPF) was
synthesized by condensing maleic anhydride with 1–2,pro-
pylene glycol as described elsewhere (Jayabalan et al., 2009).
The biosynthetic copolymer poly(propylene fumarate)-algi-
nate, AP was prepared using HT-PPF and sodium alginate.
3.7 g HT-PPF was accurately weighed and warmed to 80–90 1C
for 10 min. After cooling to 60 1C, 4 drops of conc. H
2
SO
4
were
added and stirred well. 7.4 g sodium alginate was immedi-
ately added under stirring condition and kept for more than
30 min to get poly(propylene fumarate)-alginate copolymer,
AP. The entire mixture was then slowly dissolved in 150 ml
distilled water under constant stirring and stored at room
temperature for further studies. The molecular weight of AP
was determined by HPLC using THF mobile phase at a flow
rate of 1 ml/min and polystyrene standards (Mp-100,000,
9130, and 162). The biosynthetic copolymer hydrogels were
prepared by crosslinking AP with PEGDA and vinyl monomer
at 60 1C for overnight. 60 g AP was mixed with 3.75 g
polyethylene glycol diacrylate (PEGDA) and 3.75 g each of
2-hydroxy ethyl methacrylate (HEMA). The reactants were
mixed and casted at 60 1C for overnight. The cast sheets were
treated for ionic crosslinking with Ca
2þ
and free radical
induced polymerization with ascorbic acid and ammonium
per sulfate. The hydrogel thus formed with 2-hydroxy ethyl
methacrylate (HEMA) is coded as AP-PH. Similarly, methyl
methacrylate (MMA) and N N' methylene bis acrylamide
(NMBA) were used along with PEGDA to get AP-PM and AP-
PN, respectively. The samples were cleaned and immersed in
distilled water at room temperature to leach out the
unreacted molecules. The samples were then freeze dried
overnight, sterilized by ethylene oxide and used for further
studies.
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122112
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2.3. Characterization of hydrogels
2.3.1. Physiochemical and mechanical characterization
2.3.1.1. ATR spectral studies. ATR spectrum of biosynthetic
copolymer hydrogels AP-PH, AP-PM and AP-PN were recorded
by using Nicolet 5700 FTIR Spectrometer based on ASTME
1252-98 (re-approved 2007) and E 572-01 (re-approved 2007).
2.3.1.2. Dynamic water contact angle measurement. The sur-
face properties of biosynthetic copolymer hydrogels were
carried out by dynamic contact angle studies. Six clean
samples of uniformly rectangular shape and known width
were swelled in distilled water. The advancing and receding
contact angle were determined using a Wilhelmy method on
the KSV sigma 701 tensiometer by using distilled water as a
solvent. The samples were immersed to a depth of 10 mm at
a speed of 5 mm/min ignoring the first 2 mm length. The
advancing and receding contact angle values were obtained
using the software (sg server) provided with the instrument
(Finosh et al., 2013).
2.3.1.3. Determination of equilibrium water content (EWC) and
swelling ability. The equilibrium water content (EWC) and the
swelling efficiency of hydrogels were determined by pre-
viously published procedures (Finosh and Jayabalan, 2012).
The freeze-dried samples were allowed to attain maximum
swelling in distilled water. From the dry weight and the wet
weight, EWC and swelling ability were determined.
2.3.2. Determination of surface and pore morphology
The morphology and pore size of the freeze-dried scaffolds
were determined by ESEM at low vacuum mode. The pore
average length of scaffolds was calculated from the ESEM
images using ImageJ software using multi-measure plugin.
2.3.3. Determination of crosslink density and mechanical
strength
The crosslink density of the crosslinked hydrogel materials
AP-PM and AP-PN was calculated as per the previously
published protocols (Jayabalan, 2009).The samples were taken
as circular disks and density was calculated using the volume
and the dry weight of hydrogels. The hydrogels were allowed
to attain maximum swelling in series of organic solvents
having different solubility parameters (acetone, methanol,
tetrahydrofuran, ethanol, dimethyl acetamide, toluene and
dimethyl formamide) and in water. Since water exhibited
maximum swelling, it was selected for further studies. The
freeze-dried hydrogels were allowed to attain equilibrium
swelling in water for 2 days. Both the dry weight and wet
weights were accurately measured for the determination of θ
using the following equation.
Swelling coefficient ðθÞ
¼Weight of the solvent in swollen polymer
Weight of the swelled polymer
Density of polymer
Density of solvent
From the swelling coefficient, the volume fraction Vr was
calculated. The solubility parameter of water was taken the
solubility as parameter of solvent and polymer. The crosslink
density (γ) was calculated using the modified Flory–Rehner's
equation; Vr is the volume fraction, dr is density of the
polymer, χis polymer–solvent interaction and Vo is molar
volume of the solvent.
Cross link density;γ¼
½Vr þχVr2þlnð1VrÞ
drVoð
ffiffiffiffiffiffi
Vr
3
pðVr=2ÞÞ
The tensile properties of the crosslinked hydrogel materi-
als AP-PM and AP-PN were determined using water-swollen
samples. The samples were cut to dumb bell shape using
die and tensile properties were determined as per ISO 527-2
type 5A as reported already (Finosh et al., 2013). A universal
automated mechanical test analyzer (Instron, model 3345)
connected with long travel extensometer was used.
2.4. In vitro degradation studies
2.4.1. Stability in DMEM
The stability of hydrogels in cell culture medium (DMEM) was
determined by swelling them in the medium at physiological
temperature and the solubility was noted after 7 days. Three
samples were used from each system.
2.4.2. Degradation in PBS
The long-term stability of hydrogels was determined by aging
in the PBS for 30 days as reported already (Finosh et al., 2013).
The weight loss and changes in pH, conductivity and total
dissolved solids were monitored in regular intervals of time
using the pH meter (cyber scan pc510). The initial weight was
noted. After aging, samples were removed from the medium
at an interval of 7 days and freeze-dried. The dry weight after
lyophilisation was determined. The weight loss was calcu-
lated. The changes in pH, conductivity and total dissolved
solids in the medium were monitored.
2.5. Biological evaluation
2.5.1. Effect of AP hydrogels on blood compatibility
The influence of crosslinked hydrogel materials AP-PM and
AP-PN on blood compatibility was evaluated using in vitro RBC
aggregation assay, hemolysis assay, plasma protein adsorp-
tion assay and platelet adhesion studies.
2.5.1.1. In vitro RBC aggregation assay. Red blood cells (RBC)
from healthy donor was collected with informed consent and
diluted to ten times with sterile 0.9% NaCl saline solution
after washing thrice with the same. The hydrogels were
extracted in sterile 1 PBS for 48 h. 100 μl of this PBS was
mixed with 100 μl dilute RBC suspension and incubated at
37 1C for 30 min. After incubation, cells were microscopically
examined for RBC aggregation. Positive (þve) and negative
(–ve) controls were used for comparison.
2.5.1.2. Hemolysis assay. The hemolysis induced by the
material extract to the RBC suspension was evaluated by
the standard procedures reported elsewhere (Seo et al., 2009).
2.5.1.3. Plasma protein adsorption assay. Plasma was sepa-
rated from blood by centrifuging at 1500 rpm for 10 min and
diluted to 10 ml with 1 PBS. 1 ml diluted plasma was added
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122 113
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to the PBS-swelled hydrogels and incubated for 2 h at 37 1Con
a shaker incubator. Then hydrogels were removed and the
protein fraction left out was determined by the Lowry's
method using a BSA standard (Hahm, 2011). The percentage
of proteins adsorbed onto hydrogels was quantified with
respect to the total protein content.
2.5.1.4. Platelet adhesion studies. Platelet-rich plasma (PRP)
was prepared from anticoagulated-blood by centrifuging at
2500 rpm for 5 min. Then the blood was further centrifuged at
4000 rpm for 15 min to collect PPP (platelet poor plasma). PRP
count was adjusted to 2.0–2.5 10
8
/ml with PPP, carefully
added on the center of the PBS-swelled hydrogels, and
then incubated at room temperature for 30 min in a shaker
incubator. Afterwards, hydrogels were washed 3 times with
PBS (pH-7.4) and platelets adhered on hydrogels were fixed
with 2.5% glutaraldehyde (in PBS) solution as per published
protocols (Dhandayuthapani et al., 2012). These samples were
then freeze-dried and examined in an environmental scan-
ning electron microscope.
2.5.2. In vitro evaluation of cytotoxicity
The cytotoxicity of hydrogels was evaluated by MTT assay,
the direct contact method and Live/Dead Assay.
2.5.2.1. Cell culture. The mouse fibroblast cells (L929) were
grown in DMEM supplemented with 10% FBS and penicillin
streptomycin and amphotericin-b (5000 units) in a humidified
incubator at 5% CO
2
at 3770.2 1C. The cells were regularly
monitored. The medium was changed once in three days.
The confluent monolayer was split for maintenance and
future studies.
2.5.2.2. Evaluation of the toxicity of hydrogel extracts. The
cytotoxicity of hydrogels extracts was evaluated as per ISO
10993-5 on L-929 mouse fibroblast cell culture. The extract of
the hydrogel was prepared with DMEM medium. Cell suspen-
sion containing approximately 1 10
5
cells/ml in the above
medium were seeded onto a 24 well tissue culture plate and
incubated at 37 1C and 5% CO
2
for 72 h. The percentage of the
surviving fibroblast cells were quantified by the standard
MTT assay (Idris et al., 2010).
2.5.2.3. Direct contact method. The cytotoxicity of hydrogels
under the direct contact of cell was determined by direct
contact assay. L929 fibroblast cells (1 10
4
cells/m) were
seeded on to a 12 well plate (BD Falcon) and allowed to
proliferate for 24 h to form a sub-confluent layer. Then the
hydrogel (1 cm diameter) was placed over the monolayer and
allowed to proliferate for 24 h in a CO
2
incubator. After the
incubation, cells were evaluated with respect to a control
(cells grown without hydrogels) under inverted phase con-
trast microscope attached with an imaging camera. The
images were captured using imaging software.
2.5.2.4. Live/dead assay. The L929 cells grown in contact with
the hydrogel for 5 days were evaluated with Live/Dead Assay.
A mixture of acridine orange (100 mg/ml) and ethidium bro-
mide (100 mg/ml) were added to the L929 cells and immediately
viewed under an epifluorescence microscope (Optika SRL)
using a blue filter for acridine orange and green filter for
ethidium bromide. Two images were taken from the same field
without changing settings of microscope using both filters. The
images were merged by imaging Photoshop8 CS software
(Finosh et al., 2013).
2.5.2.5. Determination of long-term viability of fibroblast by
MTT assay. The long-term viability of fibroblast adhered on
to the surface and migrated to the inner pores of all the AP
hydrogel was determined by a modified version of MTT assay
(Park et al., 2007). Around 2 10
5
cells were seeded to the
DMEM-swelled hydrogels and allowed to proliferate and
infiltrate for 3 weeks. Once on 3rd day fresh medium was
supplied to cultures and on every 6th day, hydrogels were
washed with PBS to remove the free and loosely bound cells.
Then 1 ml MTT solution (1 mg/ml) was added and incubated
at 37 1C for 3 h. After incubation, hydrogels were extracted
with the solution containing 0.01 N HCl in isopropanol. The
contents were vortexed for 30 min to detach all cells adhered
in hydrogels. It was then centrifuged at 10,000 rpm for 10 min
to settle the hydrogel particles and cell debris. Then the OD of
the supernatant was measured at 570 nm. A control was run
with cells grown without hydrogel and a reagent blank
containing scaffold without cells was run in a similar man-
ner. From the OD values the long-term viability of cells grown
inside the hydrogel were calculated.
2.6. Statistical analysis
All experiments consisted of 5 or 6 samples from each group.
The values are presented as means7standard deviations.
Statistical analysis was done with one way ANOVA using
online calculator, Statistics Calculator version-3 beta and the
level of significance was set at po0.05 for all calculations.
3. Results and discussion
3.1. Preparation and evaluation of chemically cross-linked
and mechanically favorable biosynthetic hydrogel scaffolds
The present studies involve the chemically cross-linked and
mechanically favorable biosynthetic hydrogel. The hydrogel
was prepared using the biosynthetic copolymer, poly (propy-
lene fumarate)-co-alginate copolymer, PA. The biosynthetic
copolymer PA is an oligomer with number average molecular
weight 1784 and weight average molecular weight 2010. The
addition of PEGDA to AP increases the hydrophilicity due to
its abundant hydroxyl groups and mechanical properties due
to the cross linking with the double bonds of PPF. Moreover,
the present crosslinking with PEGDA allows unreacted excess
double bonds for further crosslinking with other vinyl mono-
mers, 2-hydroxy ethyl methacrylate (HEMA), methyl metha-
crylate (MMA) and N N’methylene bis acrylamide (NMBA)
to form the chemically crosslinked and mechanically favor-
able hydrogels AP-PH, AP-PM and AP-PN, respectively. These
vinyl monomers gained importance due to their biocompat-
ibility and their ability to enhance mechanical properties
(Mitha and Jayabalan, 2009;Vallbacka and Sefton, 2007;
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122114
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Jagur-Grodzinski, 2010). The details of the synthesis are given
in Fig. 1.
The alginate used for this study is composed of 39%
guluronic acid (G) and 61% mannuronic acid (M) and pos-
sessed medium viscosity (Product no. A2033, Sigma-Aldrich,
USA). It was reported that the amount of G residues deter-
mines the mechanical properties of the alginate gels formed
by crosslinking with divalent ions, since the gelling is the
result of chelation of divalent ions with the carboxylate
anionic groups of the G residues (Drury et al., 2004). In
alginates with high M content, the degree of polymerization
is low when compared with alginates with high G content;
gels resulting from these alginates will be mechanically weak
(Becker et al., 2001). Even though gels made from high G
content alginate are mechanically stable, the inherent brit-
tleness results in the increased permeability of bioactive
molecules like antibodies (Stewart and Swaisgood, 1993;
Klein et al., 1983;Klijck et al., 1997). Moreover, the suscept-
ibility of high G content alginates to fibrosis and immunolo-
gical responses which is much higher when compared
with that of high M content alginate (Klijck et al., 1997;
De Vos et al., 1997).
Since the biocompatibility and mechanical properties are
essential for the performance of hydrogel as scaffolds for cardiac
applications, we improved the mechanical properties of high M
content alginate by copolymerizing with a mechanically robust
and biocompatible polymer, the PPF to form the AP copolymer.
The mechanical properties were further increased by reacting
with PEGDA and subsequent crosslinking with the vinyl
monomers like HEMA, MMA and NMBA as discussed in the
latter section.
Fig. 1 –Schematic representation of the synthesis of AP-P based hydrogels.
Fig. 2 –IR spectrum of freeze-dried AP-PH, AP-PM and AP-PN
hydrogels showing the effective cross linking of the vinyl
monomers.
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122 115
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The IR spectrum revealed (Fig. 2) a broad peak around
3200–3300 cm
1
for AP-PH, AP-PM and AP-PN showing the
presence of –OH groups of the alginate on the surface of the
hydrogel. With AP-PN, this peak was shifted to 3300 cm
1
indicating the stretching of both –OH and –NH groups.
With AP-PH, AP-PM and AP-PN, the peaks around 2900 cm
1
showed the asymmetric stretching of –CHgroups and
peaks around 1700 cm
1
revealed the carbonyl stretching,
both indicating the ester bond formation. The peaks around
1594–1595 cm
1
and 1412–1416 cm
1
were due to the asym-
metric and symmetric stretching of carboxylate groups indi-
cating their presence in the hydrogel surface. The peaks
around 1229–1294 cm
1
and 1024–1028 cm
1
were due to
the asymmetric and symmetric stretching of –C–O–C–of
alginate. The ATR analyses confirmed the ester bond forma-
tion and effective cross linking of the vinyl monomers.
Contact angle measurements quantify the hydrophilicity
of hydrogels. Materials having a contact angle value less than
301is considered as highly hydrophilic; if it is less than 101,
the hydrogel is super hydrophilic. The values between 301
and 901can be considered as amphiphilic. The values above
90 are indications of hydrophobicity. The hydrogels having
the contact angles ranging from highly hydrophilic to amphi-
philic range are desirable for their better blood compatibility
and cell attachment for tissue engineering applications (van
Wachem et al., 1987). The contact angles of the present
hydrogels fall in the amphiphilic range (Table 1). Therefore,
the present hydrogels will promote better cell responses and
compatibility. This amphiphilicity is due to the ample
amount of free polar functional groups like carboxylate
and hydroxyl, which is imparted especially by alginate
and PEGDA.
The swelling and EWC of a hydrogel are proportional to
the amount of water entered and imbibed; this water plays a
significant role in the diffusion of solutes from the surround-
ing medium. The micro architecture resulting from cross-
linking will also influence the water holding efficiency of the
hydrogel. Swelling properties of hydrogels are also enhanced
by the presence of ionic groups as these groups impart
additional osmotic pressure due to the presence of counter
ions. This osmotic pressure neutralizes with the swelling of
polymer chains of the hydrogel to accommodate more water
inside (Okay, 2009). The swelling and EWC for the AP-P
hydrogels are given in Table 1. The data suggest that the
AP-PM hydrogels showed maximum swelling and EWC. The
other two hydrogels AP-PH and AP-PN showed lesser but
appreciable swelling and water holding capacity. A favorable
microenvironment is orchestrated by the diluting effect of the
water present in hydrogels for cell penetration and survival.
Crosslink density was determined with respect to water
because maximum swelling was observed in water. There-
fore, the solubility product of hydrogels was close to that of
water. The results are given in Table 1. The crosslink density
is slightly higher in the case of AP-PM when compared
with AP-PN. This will be due to the increased chain length
of NMBA crosslinker than that of MMA. The Flory–Rehner
equation will not determine the chemical crosslinks exclu-
sively, but a small fraction of physical crosslinks will also be
accounted. So the obtained values were also influenced by
the physical and ionic interactions of the alginate segments.
Moreover, the chain alignment has no effect on crosslink
density by Flory-Rehner equation (Xia et al., 2013). The
comparatively less swelling and increased tensile properties
of AP-PN can be attributed to the presence of two vinyl groups
of NMBA that might have imparted higher degree of cross
linking than MMA.
Cardiac tissue consists of a well-defined extra cellular
matrix (ECM) with inherent mechanical properties, which
is crucial for the proper functioning of cardiovascular system.
Therefore, hydrogels mimicking the cardiac ECM should be
able to bear appreciable mechanical load until the repair and
regeneration is completed. The removal of entire cells from
the cardiac tissue will not significantly change the mechan-
ical properties of the cardiac ECM (Berry et al., 1975). The
evolution has universally fixed the elastic modulus of elastic
arteries including aorta of the heart within a short range of
0.3–1 MPa for most organism. This is achieved by regulating
the expression of ECM proteins (Wagenseil and Mecham,
2009). The myocardium of rat was reported to have
the stiffness of 70 kPa (Boublik et al., 2005). The stiffness
of circumferential and longitudinal right ventricle was
5478 kPa and 2074 kPa (Engelmayr et al., 2008), respectively.
The stiffness of the adult rat left ventricle was found to be
1872 kPa (Berry et al., 2006). The adult left ventricle was
found to have an ultimate tensile strength of 108 kPa with
63.8% elongation (Kallukalam et al., 2008). The tensile proper-
ties of the water swelled AP-P based hydrogels are higher
than the above-mentioned values showing their applicability
for cardiac tissue engineering (Table 2). Since these hydrogels
possess higher values than the reported ones, it can compro-
mise the loss of mechanical properties alginate fraction of the
polymer due to the exchange of monovalent ions with that of
Ca
2þ
ions. The relative higher tensile properties of AP-PN may
be due to its higher crosslinking of NMBA with respect to its
Table 1 –Surface and swelling properties of hydrogels.
Parameters (n¼6) AP-PH AP-PM AP-PN
Surface properties
Advancing angle (Po0.001) 38.8772.86 32.6772.9 33.0673.68
Receding angle (Po0.001) 39.6372.81 34.2872.94 34.0173.36
Swelling properties
EWC (Po0.001) 72.2271.33 81.4971.88 71.9575.13
% Swelling (Po0.001) 260.62717.24 444.72754.11 267.16767.58
Swelling coefficient (θ)—0.2949 0.2603
Cross-link density (γ) (g/cc) —0.1021 0.0845
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122116
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two vinyl functional groups. This is the reason for low tensile
properties and water holding ability of AP-PM, even though
the vinyl cross linker MMA is hydrophobic than NMBA.
The porosity and their morphology influence the cell
response and viability of tissue engineering scaffolds
(Bryant et al., 2007). The maximum thickness of the tissue
engineering constructs that allows the diffusion of nutrients
and oxygen was reported to be 150–200 mm. The porosity also
plays a significant role in evoking various biological
responses. The reported optimum pore size for neovascular-
ization is 5 mm, 5–15 mm for fibroblast infiltration, 100–350 mm
for bone regeneration. The micro-scale architecture can also
determine cellular orientation, clustering, penetration, and
desired function of the seeded cell (Annabi et al., 2011). If the
porosity is too small the cell penetration will be limited and
results an over confluence of cell on the seeded surface,
similar to 2D culture (Lien et al., 2009). Leor and Cohen
claimed that hydrogels for cardiac applications should have
an average porosity of 50 mm and appreciable mechanical
properties for better coordination of proper signals in a time
dependent manner (Leor and Cohen, 2004). The mechanical
properties also have implications on porosity and pore
distribution; both these properties are influenced by molecu-
lar weights of the polymer (Nwe et al., 2009). The ESEM
analysis of AP-P based hydrogels showed their characteristic
pore morphology (Fig. 3). We preferred ESEM analysis to rule
out chances of morphological variations during the analysis
due to fixing and coating procedures. Of the three, the AP-PH
had comparatively large pores as evaluated by ImageJ soft-
ware (Table 2). The greater pore length of AP-PH was due to
the hydrophilicity imparted by the HEMA crosslinker, which
promoted more water adsorption and subsequent formation
of comparatively large ice crystals during freezing. The
evacuation of these ice crystals resulted in the formation of
large pores. In all samples most of the pores remained in a
closed state. These pores are opened during the swelling
process by absorbing the medium. Due to the large pore size,
AP-PH can absorb more medium to swell resulting in rapid
degradation than the others can. However, the smaller pore
size and higher tensile strength of AP-PN makes it resistant
to degradation even though both of them possess very
similar swelling and EWC. AP-PM possessed uniform porosity
and appreciable tensile properties suited for cardiac tissue
engineering.
Table 2 –Mechanical properties of hydrogels.
Parameters (n¼6) AP-PH AP-PM AP-PN
Average pore length (Po0.001) 64.84711.54 37.9679.57 43.9377.13
Density —0.5514 0.7206
Tensile strength (KPa) (Po0.001) 10517140 6157108 16407173
Elongation at break (%) (Po0.001) 105.3278.65 79.66710.38 61.5479.87
Young modulus (KPa)(Po0.001) 20027311 12127200 603271441
Fig. 3 –ESEM analysis showing the characteristic pore morphology of AP-PH (A), AP-PM (B) and AP-PN (C) and the absence of
platelet adhesion after PRP agitation –AP-PH (D), AP-PM (E) and AP-PN (G).
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122 117
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3.2. Evaluation of degradation of hydrogels
The degradation of the AP hydrogels was evaluated in both
DMEM (containing 10% FBS) and in PBS under physiological
conditions. All hydrogels except AP-PH were stable for more
than 1 week. The DMEM-stable hydrogels, AP-PM and AP-PN
were subjected to long-term degradation studies in PBS.
AP-PM and AP-PN showed an increase in the dry weight after
1 week and 2 week of aging, respectively. This increase in the
dry weight may be due to the entrapment of some water
molecules inside the highly crosslinked networks of these
hydrogels as the initial degradation phase allows more water
particles to penetrate these hydrogels. The medium attained
a slight acidic pH. Nevertheless, in the case of AP-PN system
the pH was initially slightly acidic and propagates towards
neutrality. This slight acidity can easily be buffered by the
physiological buffers in vivo after implantation. The ionic
contents released to the medium because of degradation will
contribute to the conductivity and TDS of the medium, which
is an indication of biodegradation. The TDS and conductivity
of all hydrogels increase in a time dependent manner that
signifies the biodegradation of these hydrogels in the simu-
lated biological fluid, PBS. The data on biodegradation are
given in Fig. 4. It was reported that the rate of degradation on
the surface is constant which will maintain the bulk structure
of the hydrogel. Such degradation provides better opportu-
nities for tissue regeneration and repair (Hahm, 2011).
3.3. Evaluation of hemocompatibility
Hemocompatibility of hydrogels was assessed by RBC aggre-
gation assay, hemolytic assay, platelet adhesion assay and
indirectly by the determination of plasma protein adsorption.
The initial event occurring when blood encounters a hydrogel
is the adsorption of the plasma proteins onto the surface. The
compatibility of the material depends on the amount and
type of the protein adsorbed on the hydrogel surface (Varon
et al., 2000,Sanak et al., 2010). The surface roughness of
materials interrupts the normal blood flow leading to the
accumulation and aggregation of blood cells and subsequent
clot formation. Therefore, the surface smoothness can impart
hemocompatibility. For instance the inner lining of the blood
vessels, the endothelium, is smooth composed of cells with
less than 1 mm thickness (Finosh and Jayabalan, 2013a). RBC
membranes are susceptible to damage and rupture upon
encountering the rough surface. This leads to RBC aggrega-
tion and hemolysis. Moreover, the major hemostatic compo-
nent, the platelets, is activated upon contact with the
implanted hydrogels and lead to thrombosis. WBCs and RBCs
can also accumulate and adhere to the fibrin clot and
aggravate the immune responses (Finosh and Jayabalan,
2013b). All these events can be prevented by the adsorption
of albumin protein from the blood plasma. So the adsorbed
albumin layer on the hydrogel surface is crucial for enhan-
cing the hemocompatibility of materials (Liu et al., 2005). All
the three AP-P hydrogels are non-hemolytic (Table 3) and no
aggregation of RBCs were evident (Fig. 5). These results show
that hydrogels will not alter the membrane integrity and
function of RBCs upon contact with them. The complete
absence of platelets in the PRP treated hydrogels reveal the
potential of AP-P hydrogels to prevent the thrombus forma-
tion (Fig. 3). The plasma protein adsorption of the AP-PN
hydrogels is found to be greater when compared with that of
other two (Table 3). Since albumin is the most abundant
protein of the blood plasma, it will be the most adsorbed one
on the hydrogel surface. The SDS-PAGE analysis showed that
the band corresponding to the molecular weight of albumin
was prominent on all three hydrogels (Fig. 6). All these results
confirmed that all three AP-P hydrogels are hemocompatible
and will not evoke any adverse effects in blood flow and
compatibility.
Fig. 4 –Biodegradation profile of AP-PM and AP-PN hydrogel scaffolds.
Table 3 –Hemolysis and Protein adsorption of hydrogels.
Parameters (n¼6) AP-PH AP-PM AP-PN
% Hemolysis 0.71% 0.71% 0.71%
% Protein adsorption 4.45% 3.92% 26.7%
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122118
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3.4. Cytocompatibility of hydrogels
The cytotoxicity assays revealed the non-toxic nature of the
AP-P hydrogels. All hydrogels exhibited viability greater than
85% (Fig. 7). The AP-PH hydrogels were found to have less
stability in the culture medium. Still it promotes the viability
to 90%. These are clear indications of non-toxic nature of
degradation products or byproducts. The direct contact of
hydrogels with the sub-confluent monolayer of L929 cells did
not evoke any changes in the cell morphology and prolifera-
tion as an evident from direct contact assay (Fig. 8).
Live/dead assay revealed that all cells seeded on to the AP-
PN and AP-PM hydrogels were able to maintain their proper
health status as evident by the green fluorescence (Fig. 8). The
control cells used for live/dead assay were grown on the
tissue culture plate without hydrogels. Such cells (in 2d
environment) form a monolayer after attaining confluence
and display a uniform distribution of cells (Fig. 8F). Never-
theless, hydrogels offer a three dimensional environment to
cells due to their porous structure. Therefore, cells grown on
hydrogels seemed to be lesser in number, which can be
attributed to the increased surface area of hydrogels. It was
obvious from the ESEM analysis that the average pore length
of AP-PN was greater than that of AP-PM. This will be due to
the increase in chain length of the NMBA segments that form
the cross links with comparatively larger distance than done
by MMA in AP-PM. The cells prefer to adhere and proliferate
on the edges these pores where the nutrient availability will
be greater. This will result in the uneven distribution of cells
in hydrogels (Figs. 8D and 8E). From this, it is clear that these
Fig. 5 –RBC aggregation studies. AP-PH (A), AP-PM (B) and AP-PN (C) showing no aggregation when compared to that of PEI
treated þve control (D).
Fig. 6 –SDS-PAGE analysis of AP-PH (C), AP-PM (D) and
AP-PN (E) hydrogels after incubating with plasma showing
the prominent band corresponding to the adsorption of
albumin when compared to that of control plasma (B) and
bovine serum albumin (A).
Fig. 7 –Determination of cytotoxicity by MTT assay on
hydrogel extracts.
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122 119
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hydrogels were able to provide a favorable microenvironment
for the growth of cells.
3.5. Mechanical strength vs. long-term cell viability
in hydrogels
The mechanical properties of hydrogels can affect cell beha-
viors such as proliferation, migration, and differentiation
(Engler et al., 2006,Khatiwala et al., 2006,Pelham, Wang,
1997). The stiffness of the extracellular matrix (ECM) provides
vital instructional cues to migrating smooth muscle cells
(Engler et al., 2004,Peyton and Putnam, 2005,Wong et al.,
2003) and fibroblasts (Pelham, Wang , 1997,Lo et al., 2000) and
induce cells to migrate in a directional fashion from softer
substrates to stiffer substrate. The long-term viability of L929
cells on AP-PM and AP-PN hydrogels were quantified for 18
days (Fig. 9) and both hydrogels presented appreciable viabi-
lity even after 18 days. There was no considerable difference
between the viability. Even though the AP-PH hydrogel was
compatible, its lower mechanical strength in the culture
medium forms a hurdle for its application in the field of
cardiac tissue engineering. During the course of time, hydro-
gels will get degraded and pave way to the enlargement of
existing and opening of newer pores towards the inner side.
This will enhance the nutrient circulation, which will attract
more cells towards the interior of the scaffolds.
The mechanical properties of hydrogels should be
matched with the anatomical site of implantation. This offers
challenge for developing organ specific tissue engineering
scaffolds especially in the case of heart. Moreover, the
implanted scaffolds have to get mechanical integrity with
the host until the regeneration of the target site is completed.
Still there should be a balance between the mechanical
properties and porosity of hydrogels for the better infiltration
of cells (O'Brien, 2011). The AP-PH hydrogel was compatible
but its lower mechanical strength in the culture medium
forms a hurdle for its application in the field of cardiac
tissue engineering as a scaffold. The physiochemical,
mechanical and biological properties of the other two hydro-
gels AP-PM and AP-PN were found to be appropriate for cardiac
applications.
4. Conclusion
Mechanically robust and chemically crosslinked biosynthetic
hydrogels comprising alginate, polyproplyelene fumarate and
crosslinkers polyethylene glycol diacrylate (PEGDA) and vinyl
monomer were prepared. The effect of PEGDA and vinyl
monomers, 2-hydroxy ethyl methacrylate (HEMA), methyl
methacrylate (MMA) and N N’methylene bis acrylamide
(NMBA) on the degradation and mechanical stability and cell
viability was assessed. All three hydrogels were amphiphilic,
hemocompatible and non-cytotoxic and exhibited appreci-
able water holding capacity. Comparatively, AP-PN hydrogel
Fig. 8 –Direct contact assay of hydrogels showing the non-toxic nature of AP-PM (A) and AP-PN (B) with respect to the control
(C). The live/dead assay also revealed healthy L929 cells with AP-PM (D) and AP-PN (E) compared with that of the control (F).
Fig. 9 –Cell infiltration assay of L929 cells with AP-PM and
AP-PN hydrogels.
journal of the mechanical behavior of biomedical materials 35 (2014) 111–122120
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displayed larger pore size, increased crosslinking and higher
tensile strength, which make it resistant to degradation
even though both AP-PN and AP-PH based hydrogels possess
very similar swelling and EWC. AP-PH degraded in a faster
rate compared with that of others. Hydrogel AP-PN was
found to be mechanically strong and showed better biolo-
gical responses with long-term viability of fibroblasts viz 18
days. AP-PN hydrogel is more promising for cardiac tissue
engineering.
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