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Citation: Kamaruzaman, N.; Fauzi,
M.B.; Tabata, Y.; Yusop, S.M.
Functionalised Hybrid Collagen-
Elastin for Acellular Cutaneous
Substitute Applications. Polymers
2023,15, 1929. https://doi.org/
10.3390/polym15081929
Academic Editors: Antonia Ressler,
Inga Urlic and Alberto Romero García
Received: 2 March 2023
Revised: 10 April 2023
Accepted: 14 April 2023
Published: 18 April 2023
Copyright: © 2023 by the authors.
Licensee MDPI, Basel, Switzerland.
This article is an open access article
distributed under the terms and
conditions of the Creative Commons
Attribution (CC BY) license (https://
creativecommons.org/licenses/by/
4.0/).
polymers
Article
Functionalised Hybrid Collagen-Elastin for Acellular
Cutaneous Substitute Applications
Nurkhuzaiah Kamaruzaman 1, Mh Busra Fauzi 2, Yasuhiko Tabata 3and Salma Mohamad Yusop 1, *
1Department of Food Sciences, Faculty of Science and Technology, Universiti Kebangsaan Malaysia,
Bangi 43600, Selangor, Malaysia
2Centre for Tissue Engineering and Regenerative Medicine, Faculty of Medicine,
Universiti Kebangsaan Malaysia, Kuala Lumpur 56000, Selangor, Malaysia
3Laboratory of Biomaterials, Department of Regeneration Science and Engineering, Institute for Life and
Medical Sciences (LiMe), Kyoto University, 53 Kawara-cho Shogoin, Sakyo-Ku, Kyoto 606-8507, Japan
*Correspondence: salma_my@ukm.edu.my; Tel.: +60-132880895
Abstract:
Wound contracture, which commonly happens after wound healing, may lead to physical
distortion, including skin constriction. Therefore, the combination of collagen and elastin as the most
abundant extracellular matrix (ECM) skin matrices may provide the best candidate biomaterials for
cutaneous wound injury. This study aimed to develop a hybrid scaffold containing green natural
resources (ovine tendon collagen type-I and poultry-based elastin) for skin tissue engineering. Briefly,
freeze-drying was used to create the hybrid scaffolds, which were then crosslinked with 0.1% (w/v)
genipin (GNP). Next, the physical characteristics (pore size, porosity, swelling ratio, biodegradability
and mechanical strength) of the microstructure were assessed. Energy dispersive X-ray spectroscopy
(EDX) and Fourier transform infrared (FTIR) spectrophotometry were used for the chemical analysis.
The findings showed a uniform and interconnected porous structure with acceptable porosity (>60%)
and high-water uptake capacity (>1200%), with pore sizes ranging between 127
±
22 and
245 ±35 µm
.
The biodegradation rate of the fabricated scaffold containing 5% elastin was lower (<0.043 mg/h)
compared to the control scaffold (collagen only; 0.085 mg/h). Further analysis with EDX identified
the main elements of the scaffold: it contained carbon (C) 59.06
±
1.36–70.66
±
2.89%, nitrogen (N)
6.02
±
0.20–7.09
±
0.69% and oxygen (O) 23.79
±
0.65–32.93
±
0.98%. FTIR analysis revealed that
collagen and elastin remained in the scaffold and exhibited similar functional amides (amide A:
3316 cm−1
, amide B:
2932 cm−1
, amide I: 1649 cm
−1
, amide II: 1549 cm
−1
and amide III:
1233 cm−1
).
The combination of elastin and collagen also produced a positive effect via increased Young’s modulus
values. No toxic effect was identified, and the hybrid scaffolds significantly supported human skin
cell attachment and viability. In conclusion, the fabricated hybrid scaffolds demonstrated optimum
physicochemical and mechanical properties and may potentially be used as an acellular skin substitute
in wound management.
Keywords: elastin; collagen; hybrid bioscaffold; tissue engineering; acellular skin substitute
1. Introduction
Injuries to the skin, including chronic wounds, ulcers and burns, are among the most
prevalent and significant issues in human health care, in terms of both social and economic
impact, and they cost the US more than $28 billion per year [
1
]. Skin allografts, autografts
and xenografts can now be used to partially treat various skin injuries. However, elastin
is improperly regenerated and disorganised in scar tissue following severe burns. As
a result
, an intact elastic fibre network is absent after a cutaneous injury, contributing to
scars’ deteriorating physical properties compared with those of intact skin [
2
,
3
]. Clinicians
and researchers are realising that after a severe burn, the elastic fibre network needs to be
rebuilt for wounds to heal in a way that is both functional and aesthetically pleasing. This
makes elastin an important component of wound healing.
Polymers 2023,15, 1929. https://doi.org/10.3390/polym15081929 https://www.mdpi.com/journal/polymers
Polymers 2023,15, 1929 2 of 14
As the second most abundant protein in the human body and one of the major
components of extracellular matrix (ECM) in the skin, elastin is believed to play vital
roles in both cell signalling and structural components. Thus, its presence may lessen the
severe skin contraction and scarring linked to burn injuries while also promoting wound
healing [
4
]. Elastin may produce these effects by altering the differentiation of fibroblasts
into contractile myofibroblasts [
5
–
7
] and promoting new elastin expression in a healing
wound [
8
–
10
]. In addition, elastin has been shown to enhance angiogenesis [
11
,
12
]. Even
though elastin has demonstrated positive effects on dermal function and regeneration, its
insolubility drawbacks in three-dimensional bioscaffold construction remain underexplored.
However, given that both soluble elastin and the recombinant elastin precursor tropoelastin
are currently available, elastin-based peptides now permit a greater and more effective use
of this protein in the development of advanced wound care [4,13].
As the most abundant ECM protein, collagen is in high demand in the tissue en-
gineering, regenerative medicine and cosmetics industries. Due to its excellent biolog-
ical properties (such as low immunogenicity, porous structure, good permeability and
biodegradability), it is widely utilised in tissue engineering [
14
,
15
]. However, collagen
scaffolds often have limited strength and lack elasticity. Therefore, a dual combination of
elastin and collagen may overcome the collagen scaffold’s disadvantages of poor structural
stability and mechanical properties [
2
,
16
]. In addition, multifunctional smart biomaterial is
currently being explored in tissue engineering as an advanced technology to promote the
regeneration of skin that resembles native skin. The use of natural polymers from green
resources contributes to larger socio-economic benefits for countries worldwide. Therefore,
combining a poultry-based elastin scaffold with another natural biomaterial to replicate
the ECM microenvironment structurally, mechanically, physiologically and functionally
will be beneficial. Biomaterials derived from natural elastin are commonly extracted from
animal tissue, including bovine, porcine, poultry and marine sources [
17
–
20
]. Chicken is
considered an essential source of protein; its consumption has been quite high in recent
years, possibly because there is not much religious prohibition against chicken consumption
compared to that of other meats. Hence, poultry-based elastin represents a new generation
of elastin biomaterial widely accepted by all religious communities and cultures.
In our previous studies, elastin peptide from chicken skin has been successfully
extracted and characterised [
17
,
21
–
23
]. However, its potential as a biomaterial for skin
tissue engineering has not been explored. Therefore, in this study we developed a hybrid
scaffold composed of collagen and elastin, and then explored its physicochemical and
mechanical characteristics (in the form of a 3D-bioscaffold) to assess its potential for future
use in skin tissue engineering. In addition, both its cellular biocompatibility and its toxicity
towards human skin cells were evaluated.
2. Materials and Methods
This research was approved by the Research Ethics Committee Universiti Kebangsaan
Malaysia (UKM PPI/111/8/JEP-2020-453).
2.1. Extraction and Purification of Elastin from Broiler Skin
The method reported in our previous studies was used to extract elastin from broiler
skin [
23
,
24
]. After thawing, the broiler skin was manually defatted. The skin was then im-
mersed in NaCl overnight in a cold room. The homogenate was centrifuged at
11,000 rpm
(Eppendorf 5804R, Hamburg, Germany). Consequently, the pellet was washed and under-
went further defatting steps. The treated sample was then suspended in NaOH and heated
with constant shaking. The residues of NaOH-insoluble material were washed several
times in water and lyophilised before further analyses. Subsequently, treatment with acid
was done at 100 ◦C to enable water-soluble elastin formation before freeze-drying.
Polymers 2023,15, 1929 3 of 14
2.2. Extraction and Purification of Collagen Type-I from Ovine Tendon
Collagen extraction and purification processes were conducted according to the
method given in a previous study [
15
]. The crude tendon was cleaned of fur, skin, fas-
cia and muscle tissues. The tendon was then dissolved in 0.35 M acetic acid at 4
◦
C
for
24–48 h
until it was swollen. Subsequently, the tendon was blended and centrifuged
(5000 rpm,
5 min
, 4
◦
C). The supernatant was mixed with sodium chloride (0.05 g/mL);
(Sigma-Aldrich, St. Louis, MO, USA) and kept at 4
◦
C for 24 h. After that, it was cen-
trifuged at 5000 rpm for 5 min. The collagen pellet was then run through a dialysis tube
(
molecular weight cut-off = 14 kDa
) (Sigma-Aldrich, St. Louis, MO, USA) for 72 h, with
distilled water as the dialysis buffer being changed every 12 h. Collagen that had been
dialysed was freeze-dried for 72 h and then re-dissolved in 0.35 M acetic acid to make
a collagen concentration of 15 mg/mL for further use.
2.3. Fabrication of Collagen-Elastin (Col/Elas) Scaffolds
Col/Elas scaffolds were fabricated by combining different percentages of elastin with
the collagen solution. Briefly, the 15 mg/mL collagen solution was first pipetted into
a centrifuge tube
. Then, 2–5% of elastin powder was weighed and mixed with the collagen
solution using a mixer. After that, the solution was pipetted into the desired mould and
pre-frozen at
−
80
◦
C for 6 h, and then freeze-dried for 24–48 h. For the control group, the
scaffolds were made of collagen only (Col). The fabricated scaffolds were subsequently
cross-linked with 0.1% genipin solution (Wako, Japan) at 25
◦
C for 6 h. After 6 h of
immersion, the genipin solution was removed, and then the scaffolds were pre-frozen and
freeze-dried for 72 h. The scaffolds were as follows: Col (collagen solution only), Col/Elas-
2 (2% elastin added), Col/Elas-3 (3% elastin added), Col/Elas-4 (4% elastin added) and
Col/Elas-5 (5% elastin added).
2.4. Degree of Crosslinking
Approximately 10 mg of the test sample was weighed and heated with 200
µ
L of
10
×
ninhydrin solution (Sigma-Aldrich, St. Louis, MO, USA) for 20 min. After the test
solution was cooled to room temperature and diluted in 200
µ
L of 95% ethanol, the optical
absorbance of the solution was measured using a UV-visible spectrophotometer at 570 nm
with glycine (Gly) at various concentrations (0.006, 0.0125, 0.025, 0.05 and 0.1 mg/mL) as
the standard. The amount of free amino groups in hybrid Col/Elas scaffolds before (C
i
)
and after (C
f
) crosslinking is proportional to the optical absorbance of the solution. The
degree of crosslinking of the scaffolds was calculated using the equation below:
Degree of crosslinking (%)=Ci−Cf
Ci
×100 (1)
2.5. Swelling Ratio
First, the initial dry weight of the scaffolds was determined (W
0
). The samples were
then immersed in 2 mL of PBS (Sigma-Aldrich, St. Louis, MO, USA) for 24 h at room
temperature. The excess water was removed using filter paper and the final weight of the
sample (W1) was recorded. The swelling ratio was calculated as follows:
Ws(%)=W1−W0
W0
×100 (2)
2.6. Porosity
The initial dried weight (W
1
) of each scaffolds was measured. The scaffolds were then
immersed in absolute ethanol 99.5% (w/w) (Systerm, Shah Alam, Malaysia) overnight.
Next, the scaffolds were removed from the ethanol and blotted with filter paper to remove
excess ethanol and immediately weighed (W
2
). The porosity of the scaffolds was mea-
Polymers 2023,15, 1929 4 of 14
sured using the liquid replacement method [
25
]. Porosity (%) was calculated using the
equation below:
Porosity (%)=M2−M1
ρV×100 (3)
where V denotes the volume of the scaffolds and
ρ
refers to the density of absolute ethanol
(0.789 g/mL).
2.7. Biodegradation
The weight loss of the scaffolds after a certain period was used to calculate the
biodegradation rate of the scaffolds. The samples’ initial weight (dry) was determined
(W
0
), and they were then immersed in 2 mL of 0.0006% (w/v) collagenase (Worthington,
Lakewood, NJ, USA), which was followed by a 37
◦
C incubation. The collagenase was
removed from the samples after 24 h. After that, the samples were gently rinsed with
distilled water before being freeze-dried. Following that, the weight (W
t
) was recorded.
The weight loss rate of the scaffolds was calculated using the equation below:
Biodegradation rate(mg/h)=W0−Wt
h(4)
where W0is the dry weight, Wtis the final weight in a dry form and h is the time taken.
2.8. Chemical Characterisation of Scaffolds
Chemical characterisation of the Col/Elas scaffolds was carried out using EDX and
FTIR. Scaffolds were scanned for EDX using an INCA Energy 2000 microscope (Oxford
Instruments, Abingdon, UK) to characterise the main elements containing carbon, nitro-
gen and oxygen compositions. XRD was used to obtain atomic-scale structural infor-
mation about the scaffolds. FTIR spectra were obtained using a Perkin Elmer Spectrum
400 FTIR Spectrometer with wavenumbers ranging from 400 to 4000 cm
−1
(PerkinElmer,
Waltham, MA, USA).
2.9. Microstructure of Scaffolds
Scaffolds were prepared and visualised under the Leo 1450VP-SEM (Zeiss, Dublin, CA,
USA). The freeze-dried scaffolds were placed on the mounting platform using
two-sided
carbon tape and then coated with gold using a Polaron SC 7680 sputter-coating device
(Polaron, London, UK) for 60 s. Subsequently, the samples were introduced into the
specimen chamber and their surface morphology was examined at a 20 kV accelerating
voltage. The scaffolds were viewed at 100×magnification.
2.10. Tensile Strength
Several 3 cm
2
pieces of scaffolds were cut and attached to an Instron 8874 (Instron,
Norwood, MA, USA) using a clamp at both ends. A 50 N load transducer at a crosshead
velocity of 0.05 mm/min was used to evaluate their mechanical strength.
2.11. Human Skin Cell Harvest and Culture
Redundant skin samples were collected from all consenting healthy patients who
underwent abdominoplasty or face-lift surgery. Briefly, a 3 cm
2
skin sample was cleaned of
unwanted fragments such as fat, hair and debris, then minced into small pieces (approxi-
mately 2 mm
2
). The skin was then digested with 0.6% Collagenase Type-I (Worthington,
Lakewood, NJ, USA) for 4 h in a 37
◦
C incubator shaker, followed by cell dissociation with
0.05% Trypsin-EDTA (Gibco, Carlsbad, CA, USA) for 15 min. Digested skin containing both
keratinocytes and fibroblasts was then resuspended in co-culture medium (an equivalent
mixture of keratinocyte growth medium and fibroblast growth medium). Every 2–3 days,
the waste medium was changed.
Following the protocol established in previous studies [
15
], fibroblasts were sep-
arated from co-cultured keratinocytes using differential trypsinisation after reaching
Polymers 2023,15, 1929 5 of 14
70–80% confluence
. HDF at passage 2–3 were seeded at a density of 5000 cells/cm
2
onto
the scaffold. Cell attachment was measured after 24 h of culture. The samples were gently
washed with Dulbecco’s Phosphate Buffered Saline (DPBS) (Sigma, Livonia, MI, USA).
After that, the solution was transferred into a tube and centrifuged. The cells in DPBS were
then counted using a hemocytometer and 0.4% trypan blue solution (Sigma). The following
formula was used to calculate cell attachment:
Cell attachment (%)=Ic−Nc
Ic(5)
where Icis the initial cell seeding and Ncis the number of cells in DPBS.
2.12. Live and Dead Assay
After 24 h of seeding HDF on scaffolds, the scaffolds were stained using Invitrogen’s
LIVE/DEAD
®
Viability/Cytotoxicity Kit for mammalian cells in accordance with the
manufacturer’s instructions. A working solution containing 2
µ
M calcein AM and 4
µ
M
ethidium homodimer 1-red (EthD-1) in F-12: Dulbecco’s Modified Eagle Medium (1:1; FD;
Gibco) was prepared. Cells were washed with DPBS and then incubated with the working
solution. After 30 min, the cells were examined using a Nikon A1R confocal microscope
(Nikon, Tokyo, Japan).
2.13. Cell Morphology
Morphological features of the cells on scaffolds were examined using FESEM. Scaffolds
seeded with the cells were firstly fixed with 3% glutaraldehyde for 24 h before being
subjected to dehydration with alcohol at different percentages (50%, 75% and 95%) for
10 min
each. Then, all samples were sent for critical point drying (CPD) before viewing their
morphology using LEO 1450 VP-SEM (Zeiss, Oberkochen, Germany) at 500
×
magnification.
2.14. Cell Viability
The viability of HDF cells on the scaffolds at 1, 5, and 7 days was analysed using the
MTT assay kit according to the manufacturer’s recommendations. In each well, 100
µ
L
of fresh medium was added, followed by 10
µ
L of 12 mM MTT reagent; this was then
incubated for 4 h at 37
◦
C. The dissolution reagent was prepared by adding 10 mL of
0.01 M
HCl to one tube containing 1.0 g of SDS. Then, 100
µ
L of dissolution reagent was added
to the solution and incubated for 4 h at 37
◦
C. The absorbance was read at 565 nm. Cell
viability was calculated as follows:
Cell viability (%)=As
Ac
×100 (6)
where Asis the absorbance of the sample and Acis the absorbance of the control sample.
2.15. Statistical Analysis
The data were analysed with GraphPad Prism 8.0 (Graph Pad Software, Inc.,
San Diego, CA, USA). To compare the control with multiple groups, a one-way analy-
sis of variance (ANOVA) was used, followed by a post-hoc Tukey test. Results were
expressed as means and standard deviations. A difference was considered significant when
the p-value was <0.05.
3. Results
3.1. Physical Characterisation of Hybrid Scaffold
The gross appearances of the Col and hybrid Col/Elas scaffolds are illustrated in
Figure 1a. All the crosslinked scaffolds showed a characteristically opaque structure and
were slightly yellowish. The cross-sectional FESEM images of these scaffolds are given in
Figure 1b. According to the morphology of the samples, all scaffolds had interconnected
Polymers 2023,15, 1929 6 of 14
pores with an average pore size ranging from 100
µ
m to 245
µ
m. The pore size was also
uniformly and normally distributed across the samples.
Polymers2023,15,xFORPEERREVIEW6of15
3.Results
3.1.PhysicalCharacterisationofHybridScaffold
ThegrossappearancesoftheColandhybridCol/ElasscaffoldsareillustratedinFig-
ure1a.Allthecrosslinkedscaffoldsshowedacharacteristicallyopaquestructureandwere
slightlyyellowish.Thecross-sectionalFESEMimagesofthesescaffoldsaregiveninFigure
1b.Accordingtothemorphologyofthesamples,allscaffoldshadinterconnectedpores
withanaverageporesizerangingfrom100µmto245µm.Theporesizewasalsouni-
formlyandnormallydistributedacrossthesamples.
TheporosityofthesamplesisshowninFigure1c.Colshowedasignificantlyhigher
porosity(p<0.05)withavalueof74.83%±2.94comparedtoCol/Elas-2andCol/Elas-3,
butitwasnotsignificantlydifferentfromCol/Elas-4orCol/Elas-5.Despitethat,thepo-
rosityofthetreatedgroupsincreasedaselastinpercentageincreased.
Accordingtothedata,theaverageporesizesoftheCol/Elas-2,Col/Elas-3,Col/Elas-
4andCol/Elas-5were147±31µm,143±25µm,150±25µmand245±35µm,respectively
(Figure1d).Thesevaluesweresignificantlyhigher(p<0.05)thanthatofCol,whichwas
127±22µm.Therefore,byvaryingthecollagen/elastinratio,theporesizeofthescaffold
couldbecontrolled:thehighertheelastinpercentage,thebiggertheporesize.
(a)
(b)
(c)(d)
✱
✱
Figure 1.
(
a
) Gross appearance of the genipin-crosslinked scaffolds after crosslinking at 25
◦
C.
(b) The cross-sectional
FESEM images of the scaffolds at 100
×
magnification (Scale bar = 200
µ
m).
(c) Measured
porosity of Col/Elas scaffolds. (
d
) Mean pore sizes of Col/Elas scaffolds. (*) represents
a significant difference (p< 0.05; n= 3) between the scaffolds.
The porosity of the samples is shown in Figure 1c. Col showed a significantly higher
porosity (p< 0.05) with a value of 74.83%
±
2.94 compared to Col/Elas-2 and Col/Elas-3,
but it was not significantly different from Col/Elas-4 or Col/Elas-5. Despite that, the
porosity of the treated groups increased as elastin percentage increased.
According to the data, the average pore sizes of the Col/Elas-2, Col/Elas-3, Col/Elas-4
and Col/Elas-5 were 147
±
31
µ
m, 143
±
25
µ
m, 150
±
25
µ
m and 245
±
35
µ
m, respectively
(Figure 1d). These values were significantly higher (p< 0.05) than that of Col, which was
127
±
22
µ
m. Therefore, by varying the collagen/elastin ratio, the pore size of the scaffold
could be controlled: the higher the elastin percentage, the bigger the pore size.
Figure 2a shows the degree of crosslinking of the scaffolds. This analysis was done to
determine the crosslinking effectiveness of each sample, as assessed via a ninhydrin assay.
The results showed that the degree of crosslinking was more than 80% and significantly
Polymers 2023,15, 1929 7 of 14
increased for Col/Elas-3, Col/Elas-4 and Col/Elas-5. This indicates effective crosslinking
when compared to Col (75%). The biodegradation rates of the scaffolds, which were
determined via an enzymatic approach, are presented in Figure 2b. The biodegradation
rate of Col/Elas-5 (0.04
±
0.005 mg/h) was the slowest (p< 0.05) as compared to those of
the other scaffolds. It is predicted that 10 mg of Col/Elas-5 scaffold will be fully degraded
eight days after implantation.
Polymers2023,15,xFORPEERREVIEW7of15
Figure1.(a)Grossappearanceofthegenipin-crosslinkedscaffoldsaftercrosslinkingat25°C.(b)
Thecross-sectionalFESEMimagesofthescaffoldsat100×magnification(Scalebar=200µm).(c)
MeasuredporosityofCol/Elasscaffolds.(d)MeanporesizesofCol/Elasscaffolds.(*)representsa
significantdifference(p<0.05;n=3)betweenthescaffolds.
Figure2ashowsthedegreeofcrosslinkingofthescaffolds.Thisanalysiswasdoneto
determinethecrosslinkingeffectivenessofeachsample,asassessedviaaninhydrinassay.
Theresultsshowedthatthedegreeofcrosslinkingwasmorethan80%andsignificantly
increasedforCol/Elas-3,Col/Elas-4andCol/Elas-5.Thisindicateseffectivecrosslinking
whencomparedtoCol(75%).Thebiodegradationratesofthescaffolds,whichwerede-
terminedviaanenzymaticapproach,arepresentedinFigure2b.Thebiodegradationrate
ofCol/Elas-5(0.04±0.005mg/h)wastheslowest(p<0.05)ascomparedtothoseofthe
otherscaffolds.Itispredictedthat10mgofCol/Elas-5scaffoldwillbefullydegraded
eightdaysafterimplantation.
Theswellingratio,orwaterabsorptioncapacity,oftheCol/Elasscaffoldsisshownin
Figure2c.Theswellingratioforallsampleswasmorethan1200%.Thewaterabsorption
capacityincreasedsignificantlywiththeadditionofelastin4–5%intreatedgroupscom-
paredtoCol.TheCol/Elas-5exhibitedthehighestwaterabsorptionof1799±167%.
TheresultsofYoun g’smodulusareshowninFigure2d.Scaffoldswitha5%addition
ofelastinhadthehighestvaluesoftheYo ung’smodulus.Col/Elas-5hadoverthreetimes
theYou ng’smodulusvalueoftheColscaffold(0.22vs.0.07GPafortheCol/Elas-5and
Colscaffolds,respectively).
(a)(b)
(c)(d)
Figure2.(a)ThedegreeofcrosslinkingofCol/Elasscaffoldscrosslinkedby0.1%genipin.(b)The
enzymaticdegradationofCol/Elasscaffoldsin0.0006%collagenaseat37°C.(c)Thewaterabsorp-
tioncapacity(%)ofCol/Elasscaffolds.(d)Young’smodulusofCol/Elasscaffolds.(*)representsa
significantdifference(p<0.05;n=3)betweenfabricatedscaffolds.
Figure 2.
(
a
) The degree of crosslinking of Col/Elas scaffolds crosslinked by 0.1% genipin.
(b) The
enzymatic degradation of Col/Elas scaffolds in 0.0006% collagenase at 37
◦
C. (
c
) The water absorption
capacity (%) of Col/Elas scaffolds. (
d
) Young’s modulus of Col/Elas scaffolds. (*) represents
a significant difference (p< 0.05; n= 3) between fabricated scaffolds.
The swelling ratio, or water absorption capacity, of the Col/Elas scaffolds is shown
in Figure 2c. The swelling ratio for all samples was more than 1200%. The water absorp-
tion capacity increased significantly with the addition of elastin 4–5% in treated groups
compared to Col. The Col/Elas-5 exhibited the highest water absorption of 1799 ±167%.
The results of Young’s modulus are shown in Figure 2d. Scaffolds with a 5% addition
of elastin had the highest values of the Young’s modulus. Col/Elas-5 had over three times
the Young’s modulus value of the Col scaffold (0.22 vs. 0.07 GPa for the Col/Elas-5 and
Col scaffolds, respectively).
3.2. Chemical Characterisation of Hybrid Bioscaffold
EDX was conducted to detect the elements that existed in the samples. The identified
elements were carbon (C), oxygen (O) and nitrogen (N), which are the elements of the
amide functional groups of both collagen and elastin (NH, C=O), as shown in Table 1.
Carbon showed the highest percentage of atoms compared to other elements due to its
abundance. The addition of elastin decreased the percentages of C and N, but it increased
the percentage of O.
Polymers 2023,15, 1929 8 of 14
Table 1.
Elemental analysis by EDX in Col/Elas scaffolds. (*) represents a significant difference
(p< 0.05; n= 3) from the control group, Col.
Sample Element Weight (%)
C O N
Col 75.51 ±2.48 23.37 ±0.69 8.56 ±0.83
Col/Elas-2 70.66 ±2.89 * 23.79 ±0.65 6.70 ±0.62 *
Col/Elas-3 64.91 ±2.07 * 29.06 ±0.60 * 7.04 ±0.18 *
Col/Elas-4 72.50 ±2.00 24.07 ±0.99 6.02 ±0.20 *
Col/Elas-5 59.06 ±1.36 * 32.93 ±0.98 * 7.09 ±0.69
The FTIR spectra are presented in Figure 3: they showed typical spectra for proteins
consisting of amides A, B, I, II and III. The prominent peaks of the amide I and amide II
bands caused by the protein structures of each scaffold, were common to all samples. The
stretching bonds of C=O corresponded with amide I bands. For the examined samples,
these bands appeared at a wavelength of about 1649 cm
−1
. Amide II bands were connected
to coupled N-H bending bonds and a C-N stretching bond. These bands occurred at
about 1549 cm
−1
. Both collagen and elastin from the porcine aorta have similar bands of
amide I (approximately 1580–1718 cm
−1
) and amide II (approximately 1480–1580 cm
−1
)
absorbances [
26
,
27
]. Amide A bands, which represent stretching N-H bonds, occurred at
approximately 3316 cm
−1
. The stretching vibrations of N-H in the Fermi resonance, which
occurred at around 2932 cm
−1
, formed amide B bands. Another band was observed at
a peak
of 1233 cm
−1
for amide III from the CH
2
wagging vibration of proline side chains,
which could have originated from the similar proline content in collagen and elastin [
26
,
28
].
Polymers2023,15,xFORPEERREVIEW8of15
3.2.ChemicalCharacterisationofHybridBioscaffold
EDXwasconductedtodetecttheelementsthatexistedinthesamples.Theidentified
elementswerecarbon(C),oxygen(O)andnitrogen(N),whicharetheelementsofthe
amidefunctionalgroupsofbothcollagenandelastin(NH,C=O),asshowninTable1.
Carbonshowedthehighestpercentageofatomscomparedtootherelementsduetoits
abundance.TheadditionofelastindecreasedthepercentagesofCandN,butitincreased
thepercentageofO.
Tab le 1.ElementalanalysisbyEDXinCol/Elasscaffolds.(*)representsasignificantdifference(p<
0.05;n=3)fromthecontrolgroup,Col.
SampleElementWei gh t(%)
CON
Col75.51±2.4823.37±0.698.56±0.83
Col/Elas-270.66±2.89*23.79±0.656.70±0.62*
Col/Elas-364.91±2.07*29.06±0.60*7.04±0.18*
Col/Elas-472.50±2.0024.07±0.996.02±0.20*
Col/Elas-559.06±1.36*32.93±0.98*7.09±0.69
TheFTIRspectraarepresentedinFigure3:theyshowedtypicalspectraforproteins
consistingofamidesA,B,I,IIandIII.TheprominentpeaksoftheamideIandamideII
bandscausedbytheproteinstructuresofeachscaffold,werecommontoallsamples.The
stretchingbondsofC=OcorrespondedwithamideIbands.Fortheexaminedsamples,
thesebandsappearedatawavelengthofabout1649cm−1.AmideIIbandswereconnected
tocoupledN-HbendingbondsandaC-Nstretchingbond.Thesebandsoccurredatabout
1549cm−1.BothcollagenandelastinfromtheporcineaortahavesimilarbandsofamideI
(approximately1580–1718cm−1)andamideII(approximately1480–1580cm−1)absorb-
ances[26,27].AmideAbands,whichrepresentstretchingN-Hbonds,occurredatapprox-
imately3316cm−1.ThestretchingvibrationsofN-HintheFermiresonance,whichoc-
curredataround2932cm−1,formedamideBbands.Anotherbandwasobservedatapeak
of1233cm−1foramideIIIfromtheCH2waggingvibrationofprolinesidechains,which
couldhaveoriginatedfromthesimilarprolinecontentincollagenandelastin[26,28].
Figure3.TheFTIRspectraofCol/Elasscaffoldscrosslinkedby0.1%genipin.
Amide I (1649 cm−1)
Amide B (2932 cm−1)Amide III (1233 cm−1)
Amide A (3316 cm−1)Amide II (1549 cm−1)
Wavenumbers (cm
-1
)
01000200030004000
% Transmittance
Col/Elas-2
Col/Elas-3
Col/Elas-4
Col/Elas-5
Col
Figure 3. The FTIR spectra of Col/Elas scaffolds crosslinked by 0.1% genipin.
3.3. Biocompatibility of Elas/Col Scaffolds
It was found that HDF cultured on all scaffolds were viable (Figure 4a), as the presence
of red HDF on scaffolds were few (white circles), and green HDF predominated, indicating
no cytotoxic effect. The percentages of dead dells for Col, Col/Elas-2, Col/Elas-3, Col/Elas-
4 and Col/Elas-5 were 16.2%, 1.6%, 3.5%, 9.2% and 16.6%, respectively. In comparison
to the other samples, the cells on Col/Elas-2 had a round shape and less spindles. The
samples showed good cell attachment within 24 h, where more than 80% of the cells suc-
cessfully adhered onto the scaffolds as shown in Figure 4b. However, no differences in HDF
attachment were observed between the Col scaffold and the other scaffolds. The FESEM
images of the HDF morphologies on the fabricated scaffolds are shown in Figure 4d. Briefly,
all scaffolds promoted cell adhesion and appeared round that had an average diameter
Polymers 2023,15, 1929 9 of 14
of 20
µ
m with a surrounding matrix. They were found inside the scaffold, indicating that
the cells migrated and attached successfully. Further evaluation was performed on the cell
viability of HDF scaffolds using MTT assays at 1, 5 and 7 days (Figure 4c). Generally, each
scaffold showed increased cell viability from day 1 to day 7. However, the elastin-based
scaffolds (Col/Elas-3 and Col/Elas-4) showed less cell viability on day 5 as compared to
Col, but they had higher viability on day 7.
Polymers2023,15,xFORPEERREVIEW9of15
3.3.BiocompatibilityofElas/ColScaffolds
ItwasfoundthatHDFculturedonallscaffoldswereviable(Figure4a),asthepres-
enceofredHDFonscaffoldswerefew(whitecircles),andgreenHDFpredominated,in-
dicatingnocytotoxiceffect.ThepercentagesofdeaddellsforCol,Col/Elas-2,Col/Elas-3,
Col/Elas-4andCol/Elas-5were16.2%,1.6%,3.5%,9.2%and16.6%,respectively.Incom-
parisontotheothersamples,thecellsonCol/Elas-2hadaroundshapeandlessspindles.
Thesamplesshowedgoodcellaachmentwithin24h,wheremorethan80%ofthecells
successfullyadheredontothescaffoldsasshowninFigure4b.However,nodifferencesin
HDFaachmentwereobservedbetweentheColscaffoldandtheotherscaffolds.The
FESEMimagesoftheHDFmorphologiesonthefabricatedscaffoldsareshowninFigure
4d.Briefly,allscaffoldspromotedcelladhesionandappearedroundthathadanaverage
diameterof20μmwithasurroundingmatrix.Theywerefoundinsidethescaffold,indi-
catingthatthecellsmigratedandaachedsuccessfully.Furtherevaluationwasper-
formedonthecellviabilityofHDFscaffoldsusingMTTassaysat1,5and7days(Figure
4c).Generally,eachscaffoldshowedincreasedcellviabilityfromday1today7.However,
theelastin-basedscaffolds(Col/Elas-3andCol/Elas-4)showedlesscellviabilityonday5
ascomparedtoCol,buttheyhadhigherviabilityonday7.
(a)
(b)
Polymers2023,15,xFORPEERREVIEW10of15
(c)
(d)
Figure4.Cell-scaffoldinteractionin3Dscaffolds.(a)Live/Deadassayofthefabricatedscaffolds;
Scalebar:100µm.(b)Thequantificationofcellaachmentat24h.(c)CellviabilityofHDFonthe
scaffoldsatday1,5and7quantifiedbyMTTassay.(d)
Cross-sectionalmorphologicalfeaturesof
HDFat500×magnification;Scalebar:20µm.(*)representsasignificantdifference(p<0.05;n=3)
betweenfabricatedscaffolds.
4.Discussion
Multifunctionalsmartbiomaterialiscurrentlybeingexploredintissueengineering
asanadvancedtechnologythatpromotestheregenerationofskinthatresemblesnative
skin.However,whencreatingorestablishingthesuitabilityofascaffoldforuseintissue
engineering,severalcrucialfactorsmustbeconsidered,includingbiocompatibility,bio-
degradability,mechanicalproperties,porosityandswellingratio,allofwhichsupport
tissueregeneration[29,30].
Thus,theaimofthisstudywastofabricatehybridscaffoldscontainingelastinand
collageninordertoevaluatetheeffectofaddingelastinonthescaffold’sperformancein
optimisinghealingcapacitywhilemaintainingabeeroverallappearance.Grossappear-
ancewasexaminedtoevaluatethephysicalappearanceofthefabricatedscaffolds.The
resultsshowedthattheadditionof2–5%ofelastinpowderdidnotchangethecolourof
thescaffoldwhencomparedtoaColsample.Allsamplesunderwentpost-crosslinking
with0.1%(w/v)genipintoalterthefinalstructural,mechanicalandbiologicalcharacter-
isticsofthescaffolds.Anaturalcrosslinker,genipinismadefromthefruitsoftheGardenia
jasminoidesplant.Itisnon-toxicanddegradesmoreslowlythanchemicalcrosslinkerssuch
asglutaraldehyde[31,32].Somesizeshrinkagewasobservedafterthescaffoldswere
crosslinked.Crosslinkingcausesthesamplestoformintra-andintermolecularcrosslink-
ingnetworkswiththefreeaminogroupsofthesamples[33],therebyleadingtosize
shrinkages.
Inaddition,bothcollagenandelastinareproteins;therefore,increasingtheelastin
concentrationmayformadditionalsubstrateforthecrosslinkingprocessandincreasethe
crosslinkingefficiencyofthescaffolds[34].Boththetemperatureandgenipinconcentra-
tionusedduringcrosslinkingarecrucialfactorsthatcanaffecttheefficiencyofcrosslink-
ing.Itwasdiscoveredthattheidealgenipinconcentrationforagelatinescaffoldis0.5%,
Figure 4.
Cell-scaffold interaction in 3D scaffolds. (
a
) Live/Dead assay of the fabricated scaffolds;
Scale bar: 100
µ
m. (
b
) The quantification of cell attachment at 24 h. (
c
) Cell viability of HDF on the
scaffolds at day 1, 5 and 7 quantified by MTT assay. (
d
) Cross-sectional morphological features of
HDF at 500
×
magnification; Scale bar: 20
µ
m. (*) represents a significant difference (p< 0.05; n= 3)
between fabricated scaffolds.
Polymers 2023,15, 1929 10 of 14
4. Discussion
Multifunctional smart biomaterial is currently being explored in tissue engineering
as an advanced technology that promotes the regeneration of skin that resembles native
skin. However, when creating or establishing the suitability of a scaffold for use in tis-
sue engineering, several crucial factors must be considered, including biocompatibility,
biodegradability, mechanical properties, porosity and swelling ratio, all of which support
tissue regeneration [29,30].
Thus, the aim of this study was to fabricate hybrid scaffolds containing elastin and
collagen in order to evaluate the effect of adding elastin on the scaffold’s performance in op-
timising healing capacity while maintaining a better overall appearance. Gross appearance
was examined to evaluate the physical appearance of the fabricated scaffolds. The results
showed that the addition of 2–5% of elastin powder did not change the colour of the scaffold
when compared to a Col sample. All samples underwent post-crosslinking with
0.1% (w/v)
genipin to alter the final structural, mechanical and biological characteristics of the scaf-
folds. A natural crosslinker, genipin is made from the fruits of the
Gardenia jasminoides
plant. It is non-toxic and degrades more slowly than chemical crosslinkers such as glu-
taraldehyde [
31
,
32
]. Some size shrinkage was observed after the scaffolds were crosslinked.
Crosslinking causes the samples to form intra- and intermolecular crosslinking networks
with the free amino groups of the samples [33], thereby leading to size shrinkages.
In addition, both collagen and elastin are proteins; therefore, increasing the elastin
concentration may form additional substrate for the crosslinking process and increase the
crosslinking efficiency of the scaffolds [
34
]. Both the temperature and genipin concentration
used during crosslinking are crucial factors that can affect the efficiency of crosslinking. It
was discovered that the ideal genipin concentration for a gelatine scaffold is 0.5%, and that
the ideal crosslinking temperature is 25
◦
C [
32
]. In this study, a slightly lower concentration
of genipin (0.1%) but the same incubation temperature were used. Therefore, further
studies could vary the genipin concentrations and time of incubation to improve the
degree of crosslinking to more than 90%. In addition, crosslinking also affects swelling and
degradation, hence the release of integrated bioactive chemicals from scaffolds. Therefore,
having optimum crosslinking while fabricating biomaterials is crucial.
To maintain a moist environment, an optimal biomaterial in cutaneous wound healing
must have a high-water absorption capacity. In this study, the fabricated hybrid scaffolds
had good water absorption ability, which aids in the prevention of wound exudate accumu-
lation in chronic skin wounds and the scaffolds expand upon contact with the wound. In
addition to that, moist wounds are known to heal faster with an adequate supply of growth
factors and other molecules to the healing tissues [
35
]. A moist environment also aids in
providing essential nutrients to cells [
35
–
37
]. Compared to a dry environment,
a moist
environment speeds up the time a wound takes to heal because it makes it easier for growth
factors to be released and for cells to multiply. It also helps basal epidermal cells to move
into the wound. In addition, a moist wound environment minimises the risk of infection by
promoting angiogenesis, lowering pH and making the wound unfriendly to bacteria [38].
The presence of hydroxyl (OH) and amino functional groups (NH
2
) of collagen and
elastin, which can interact with water via hydrogen bonds, may explain the high swelling
tendency of these scaffolds. Generally, the more crosslinking, the less swelling that is
attainable [
39
,
40
]. However, these swelling results contradicted the usual notion that
a higher
degree of crosslinking leads to a lower swelling ratio. The presence of elastin
increased the hydrophilic groups of the scaffolds, increasing their swelling ratio [
41
]. It
is well known that collagen has good absorption properties that help to keep wounds
hydrated [
42
]. Therefore, a combination of elastin and collagen could be useful for wound
healing. In addition, the porous structure of scaffolds containing elastin allows fabricated
scaffolds to entrap more water.
In the current study, all samples showed porosity of more than 60%, which is within the
range of an ideal wound dressing requirement of 60–90% porosity [
43
]. The good porous
structure of the fabricated scaffolds indicates that the freeze-drying method was both
Polymers 2023,15, 1929 11 of 14
appropriate and the simplest approach for the fabrication of 3D scaffolds in this study. The
collagen/elastin scaffolds cool to the freezing point during the freeze-drying process used to
create these scaffolds, whereupon the ice phase separates to create
a porous
structure [
44
].
Specifically, Col scaffolds displayed a more compact, sheet-like microstructure, while
treatment groups with elastin showed a fibre-like structures.
Mass transport, cell migration and tissue ingrowth into the implant’s structure are all
greatly influenced by a scaffolds’ porosity, pore size and connection types [
45
]. The ideal
pore diameter for scaffolds is thought to be between 100 and 500
µ
m [
46
]. Additionally,
pores between 90 and 130
µ
m are the ideal size for skin regeneration and are sufficient to
allow fibroblast migration and proliferation [
47
,
48
]. The collagen/elastin scaffold has good
potential for use as a skin substitute due to its interconnected pores that are crucial for
tissue repair, as well as having the right porosity and pore size.
In acellular skin substitutes, the biodegradation process is critical. Frequent dressing
changes are not cost-effective and could be a burden to both patients and caregivers. Rapid
degradation of a scaffold can impair its mechanical properties, while delayed degradation
can cause an inflammatory reaction and lead to the failure of skin implantation [
48
]. The
ideal scaffold would reduce inflammation, release medication and aid tissue regeneration
without requiring frequent replacement. Since a scaffold is just a template and not perma-
nent, it must be fully biodegradable to allow cells to produce their own ECM during the
healing process [
49
]. In this study, the addition of elastin decreased the degradation rate of
the scaffolds. A study done by Su, Fujiwara and Bumgardner [
50
] also showed that scaf-
folds with a high percentage of elastin had a slower degradation rate as compared to those
with a low percentage. In addition, the degradation results showed that highly crosslinked
samples will degrade more slowly compared with less crosslinked samples. This result
shows that crosslinking can also improve the resistance of scaffolds against enzymatic
degradation [44,50], thus enhancing the property of the scaffold as a skin substitute.
Pure collagen scaffolds have been shown to have poor mechanical characteristics.
A scaffold
needs to be strong enough to allow for surgical handling during implantation.
To address this limitation, intermolecular crosslinking was attempted to enhance the
mechanical strength and stability of the collagen scaffold [
51
]. The addition of elastin
increases Young’s modulus’ values, which are indicative of the elasticity and rigidity of the
samples [
52
]. In addition, the porosity of a scaffold is an important factor in influencing its
mechanical properties. In this study, the addition of 2–3% elastin increased both the porosity
of the scaffolds and their mechanical strength. Moreover, the addition of genipin makes the
scaffolds more rigid without changing their porosity; it also has low toxicity [
53
]. In a study
by Skopinska-Wisniewska et al. [
27
], the same pattern was observed in a high-compression
modulus in relation to all of the crosslinked collagen–elastin-based samples.
The FTIR spectra obtained in this study revealed no changes in molecular structure
after adding a mixture of collagen and elastin at various concentrations. This observation
was also supported in earlier studies [
26
,
27
,
54
], which suggested that the addition of
elastin and a crosslinking agent does not significantly affect collagen structure. Thus, it
appeared that a hybrid scaffold was successfully fabricated while still maintaining the
original characteristics from the functional groups.
Another consideration for an ideal scaffold for treating skin wounds is cellular com-
patibility, which can help human skin cells proliferate and remain viable. The results of
a Live/Dead
assay demonstrated that fibroblasts had excellent viability when adhering to
and growing on the surfaces of collagen and elastin scaffolds. This implied that these scaf-
folds were not toxic to cells and were in fact inducing cell growth. A study on aortic elastin
and collagen scaffolds also showed that fibroblasts adhered to and proliferated on scaffold
surfaces with excellent cell viability [
55
]. Even though some dead cells were observed in
this study, the number of dead cells was low compared to that of live cells. Furthermore,
the round-shaped cells observed on Col/Elas-2 in the Live/Dead assay suggested that
the cells were not fully proliferating yet, as evidenced by both the cell viability and cell
attachment results, where Col/Elas-2 viability and attachment were low after 24 h. While
Polymers 2023,15, 1929 12 of 14
there was some reduction in cell viability on day 5 for Col/Elas-3 and Col/Elas-4, it was not
significantly different from day 1. When compared to Col, however, both Col/Elas-3 and
Col/Elas-4 demonstrated significantly lower viability, most likely due to low cell prolifera-
tion rates and the need for time to become compatible with the scaffolds before accelerating
proliferation as observed on day 7 [
56
]. Additionally, it has been demonstrated that com-
bining elastin with collagen and gelatine encourages fibroblast proliferation
in vitro
[
57
].
All the abovementioned properties could lead to the better compatibility and attachment of
the scaffolds.
5. Conclusions
In this study, a collagen/elastin scaffold was successfully fabricated using collagen
from sheep and elastin from poultry, which is a low-cost, abundant and renewable material.
This study presented the effects of collagen/elastin on the physicochemical and mechanical
properties of the scaffold. In addition to its large surface area (due to a well-connected
pore network structure), the collagen scaffold with elastin exhibited higher crosslinking
efficiency (from 74.70% to 88.77%), higher water absorption (from 1012.70% to 1799.06%),
lower degradation rates (from 0.085 mg/h to 0.043 mg/h) and higher mechanical strength
(from 0.07 GPa to 0.22 GPa) than the Col scaffold. In addition, the hybrid Col/Elas scaffold
exhibited good biocompatibility with all functional groups of the individual constituents
(elastin and collagen) remaining intact and homogenous, as observed from the SEM, FTIR
and EDX results. In summary, this study demonstrated the potential of combining collagen
from ovine and poultry sources to create a hybrid scaffold with biomedical applications
as an acellular skin substitute. However, further
in vivo
studies should be performed to
better understand the potential of fabricated scaffolds for wound healing.
Author Contributions:
Methodology, N.K.; formal analysis, N.K.; investigation, N.K.; resources,
M.B.F. and S.M.Y.; writing—original draft, N.K.; writing—review & editing, M.B.F., Y.T. and S.M.Y.;
supervision, M.B.F., Y.T. and S.M.Y.; funding acquisition, M.B.F. and S.M.Y. All authors have read and
agreed to the published version of the manuscript.
Funding:
This research was funded by the Ministry of Higher Education, grant number FRGS/1/2018/
STG05/UKM/02/8.
Institutional Review Board Statement:
This research was approved by the Research Ethics Commit-
tee Universiti Kebangsaan Malaysia (UKM PPI/111/8/JEP-2020-453).
Data Availability Statement: Not applicable.
Conflicts of Interest: The authors declare no conflict of interest.
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