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Łopianiaketal.
Journal of Biological Engineering (2023) 17:20
https://doi.org/10.1186/s13036-023-00337-9
RESEARCH
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Open Access
Journal of
Biological Engineering
Multilayered blow-spun vascular prostheses
withluminal surfaces inNano/Micro range:
theinuence onendothelial cell andplatelet
adhesion
Iwona Łopianiak1,2, Wiktoria Rzempołuch1, Mehtap Civelek3, Iwona Cicha3, Tomasz Ciach1,4 and
Beata A. Butruk‑Raszeja1*
Abstract
Background In this study, two types of polyurethane‑based cylindrical multilayered grafts with internal diam‑
eters ≤ 6 mm were produced by the solution blow spinning (SBS) method. The main aim was to create layered‑wall
prostheses differing in their luminal surface morphology. Changing the SBS process parameters, i.e. working distance,
rotational speed, volume, and concentration of the polymer solution allowed to obtain structures with the required
morphologies. The first type of prostheses, termed Nano, possessed nanofibrous luminal surface, and the second
type, Micro, presented morphologically diverse luminal surface, with both solid and microfibrous areas.
Results The results of mechanical tests confirmed that designed prostheses had high flexibility (Young’s modulus
value of about 2.5 MPa) and good tensile strength (maximum axial load value of about 60 N), which meet the require‑
ments for vascular prostheses. The influence of the luminal surface morphology on platelet adhesion and the attach‑
ment of endothelial cells was investigated. Both surfaces did not cause hemolysis in contact with blood, the percent‑
age of platelet‑occupied area for Nano and Micro surfaces was comparable to reference polytetrafluoroethylene
(PTFE) surface. However, the change in morphology of surface‑adhered platelets between Nano and Micro surfaces
was visible, which might suggest differences in their activation level. Endothelial coverage after 1, 3, and 7 days of
culture on flat samples (2D model) was higher on Nano prostheses as compared with Micro scaffolds. However, this
effect was not seen in 3D culture, where cylindrical prostheses were colonized using magnetic seeding method.
Conclusions We conclude the produced scaffolds meet the material and mechanical requirements for vascular pros‑
theses. However, changing the morphology without changing the chemical modification of the luminal surface is not
sufficient to achieve the appropriate effectiveness of endothelialization in the 3D model.
Keywords Multilayered small‑diameter vascular grafts, Hemocompatibility, Endothelial cells, Solution blow spinning,
Nanofibers, Magnetic seeding
*Correspondence:
Beata A. Butruk‑Raszeja
Beata.Raszeja@pw.edu.pl
Full list of author information is available at the end of the article
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Page 2 of 17
Łopianiaketal. Journal of Biological Engineering (2023) 17:20
Introduction
Cardiovascular diseases (CVD) were responsible for 32%
of all deaths worldwide in 2019 [1]. In advanced stages
of CVD,the only choice is the surgical intervention, in
which damaged arteries are replaced with autologous
vessels or synthetic prostheses. However, the clinical
success of this procedure is limited by low availability of
autologous blood vessels. Also, commercially available
synthetic grafts made of expanded polytetrafluoroeth-
ylene (ePTFE) or polyethylene terephthalate (PET) with
diameters ≤ 6mm frequently fail. e poor patency rate
of synthetic prostheses [2–5] compels researchers to look
for new approaches and solutions.
Although intimal hyperplasia or inflammatory com-
plications may negatively affect the patency of the arti-
ficial vessels upon implantation, the main reason for
prosthetic graft failure is occlusion caused by throm-
bosis [2–5]. In physiological conditions, the lumen of
the blood vessel is covered with endothelial cells (ECs),
which actively counteract the processes of platelet aggre-
gation and blood coagulation through the synthesis and
secretion of various bioactive substances e.g.: nitric
oxide, heparan sulphate, prostacyclin [6]. In addition,
intact endothelial monolayer inhibits the proliferation
of smooth muscle cells (SMCs), limiting the risk of inti-
mal hyperplasia. Implantation of synthetic grafts with-
out this endothelium barrier may lead to surface protein
adsorption followed by platelet adhesion, activation,
and aggregation. Several strategies have been proposed
towards quick endothelialization of prosthesis’ lumi-
nal surface. One of them is invitro endothelialization,
i.e. colonization of the prosthesis with the patient’s cells
before the implantation procedure. Another approach,
insitu endothelialization, is based on colonization with
ECs in the patient’s body, which is possible through
transanastomotic growth, transmural infiltration, and
endothelialization with endothelial progenitor cells cir-
culating in bloodstream [7].
Regardless of the approach chosen, the surface of the
prosthesis must enhance the adhesion and proliferation
of ECs, to enable restoration of a functional endothelium
and, as a result, to reduce clotting processes. e litera-
ture proposes various strategies to improve EC attach-
ment. One of them is based on the modulation of the
surface topography e.g. by adding nanostructures to the
lumen surface. is strategy assumes that introduction of
nanostructures, e.g. nanofibres, increases surface to vol-
ume ratio and provides more binding sites for cell adhe-
sion and biomolecule adsorption [8].
e ideal small-diameter vascular grafts (with diam-
eter ≤ 6mm) should mimic the layered structure of the
native blood vessels and exhibit comparable mechanical
properties. is leads to the idea of a layered prosthesis,
where the inner surface is designed to provide an envi-
ronment and topography suitable for reconstructing
the endothelial layer, whereas the outer layers are tai-
lored to fulfill other, specific purposes, i.e. ensuring
appropriate mechanical properties and suitable poros-
ity to enable the ingrowth of capillaries. To date, elec-
trospinning has been the most universal and popular
method of manufacturing fibrous vascular prostheses
[9]. is technique enables the production of prosthe-
ses containing both micro- and nanofibers, as well as
layered prostheses containing fibers of various sizes
[10]. Nonetheless, electrospinning has a number of lim-
itations related to high voltage requirements, low pro-
duction rate, and limited number of suitable solvents
[11]. Our group has developed an alternative technique
for fabrication of fibrous vascular prosthesis, namely
the solution blow spinning (SBS) method [12, 13]. e
SBS system is similar to the electrospinning system
but does not require the presence of anelectric field.
e driving force of the process is the pressure of the
working gas, which is fed to the nozzle together with
polymer solution. e pressure forms fibers at the out-
let of the nozzle and deposit them on the rotating col-
lector. SBS has several advantages over electrospinning,
including low cost, easiness to scale and controlof the
parameters, as well as no need for high voltage [14, 15].
Vascular grafts can be made of natural or synthetic
polymers. Of the synthetic materials, PET or ePTFE
were originally used. ese materials are still the most
commonly used in clinical practice for peripheral vessel
replacement, but they have some disadvantages. Most
importantly, their surface does not promote cell adhesion
and they are quite resistant to chemical modifications.
is is a significant disadvantage because the majority of
synthetic polymers require surface modification in order
to improve their biological properties. PUsare a chemi-
cally diverse group of polymers that are often studied in
the context of biomedical applications. Heart valve, car-
tilage, skin, blood vessel and bone scaffolds have been
successfully produced from PUs [16–23]. e versatil-
ity of PU-based scaffolds arises from material bio- and
hemocompatibility, its easy processing, and appropriate
mechanical properties [24]. Furthermore, the mechani-
cal properties of PUs, including elasticity, strength, hard-
ness, and resiliency, are easily controllable by changing
the ratio of soft and hard segments. [25, 26]. Moreover,
PUs present attractive biological properties probably due
to the fact that urethane bond is similar to the peptide
bond. Cells, including ECs, are able to adhere to the PU
surface, even without the application of chemical modi-
fications. is makes it possible to study the influence of
the topography (fiber diameter etc.) on the cell-surface
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Page 3 of 17
Łopianiaketal. Journal of Biological Engineering (2023) 17:20
interactions. In this study, medical grade Chrono-
Flex PU was selected because of its reportedly high
athrombogenicity.
In our previous study, we evaluated the influence of
fibrous surface morphology on endothelial and smooth
muscle cell (SMC) growth [27]. We have shown that both
morphology (solid versus fibrous) and average fiber diam-
eter (submicron fibers versus microfibers) of scaffolds
influenced the growth of ECs. Here, we designed layered
cylindrical prostheses that differ in the morphology of
the luminal surface. e aim of the present work was to
compare two types of prostheses with multilayered walls.
e outer layer is made of aligned microfibers, with an
average diameter of about 1000µm, which are intended
to support the SMCs development. e middle layer with
total layer thickness of about 500µm, containing non-
aligned microfibers with an average diameter of about
1000 µm is expected to give the prosthesis adequate
flexibility and mechanical strength. Finally, the internal
layer is composed of dense microfibers presenting with
two morphological types of luminal surface. is layer is
designed to support the attachment of ECsby ensuring
the appropriate topography, either a mixed solid/micro-
fiber structure or a nanofiber structure. e prosthesis
termed “Micro” has a luminal surface composed of solid
areas (flat, film-like surfaces without fibrous structures)
and microfibers, while in the prosthesis termed “Nano”
the luminalsurface is composed of nanofibers. Following
the fabrication of the prostheses, their physical proper-
ties were characterized. Further, hemocompatibility of
the distinct luminal morphologies was compared using
human platelets, and two cell seeding models were used
to evaluate the growth of ECs on Nano versus Micro
surfaces.
Materials andmethods
Vascular prostheses fabrication
Prostheses were produced from medical grade polyu-
rethane solution by SBS method, as described else-
where [12, 27]. Briefly, polyurethane ChronoFlex®C75A
(Advanced Biomaterials, USA) was dissolved overnight
in 1,1,1,3,3,3-hexafluoro-2-propanol (> 99%Fluorochem
Ltd, UK) on magnetic stirrer. e polymer solution was
transferred into syringe and fed to the inner nozzle of
concentric nozzle system. e polymer solution flow rate
was controlled by syringe pump. e inner diameters
of inner and outer nozzles were 1.1 and 10mm, respec-
tively. Fibers were collected on rotating collector, 6mm
in diameter and 12 cm in length, mounted 10–30 cm
away from the tip of inner nozzle. Prior to the SBS pro-
cess, the collector was covered with a thin layer of 10%
w/v poly(ethylene) glycol 2000 (Sigma Aldrich, Germany)
solution in distilled water in order to simplify removal
of the prosthesis from the collector surface. After the
prosthesis deposition and its immersion (together with
the collector) in distilled water for 2min, the prosthesis
was gently slid off the collector. e slight shrinking of
the prostheses after the removal resulted in a final inner
diameter of 5mm.
Two variants of layered prostheses were produced:
(a) Nano and (b) Micro. As shown in Fig. 1A, Nano
prosthesis consists of the following layers: nanofibers
(luminal), dense microfibers, microfibers, and aligned
microfibers (outermost). Micro prosthesis consists of the
following layers: dense microfibers (luminal), microfib-
ers, and aligned microfibers (outermost). e SBS pro-
cess parameters used for producing individual layers are
shown in Table1.
Morphology oftheprostheses
e prostheses were cut open and flat samples with
dimensions 0.5 × 0.5 cm were glued to the SEM stubs
with conductive carbon adhesive tape. Samples of inter-
nal (n = 3) and external surfaces (n = 3) were prepared.
To characterize cross-sectional sample’s morphology,
samples of prostheses 0.5 cm in length (n = 3 for each
type) were glued upright to SEM stubs. e samples were
then coated with 15nm of gold using sputter coater (K550
Emitech, Quorum Technologies). Images of every sam-
ple (n = 10) were taken at magnifications × 200, × 600,
and × 5000 using scanning electron microscopy Phenom
G1 (Phenom World). SEM images were used to deter-
mine fiber diameter, pore size, and prostheses thickness.
To determine the percentage of fibrous area on the inter-
nal luminal surface of Micro prostheses, the percentage
of fibrous surface was measured in n = 20 SEM images.
In every sample, n = 100 fiber diameters were measured
using Fiji software. For nanofibrous internal surface of
Nano prostheses, pore size was determined using SEM
images of luminal surface at magnification × 5000. For
this, the threshold tool (Fiji software) was used to deline-
ate the most surface pores and the area of n = 100 pores
was measured using Fiji software e pores were approx-
imated to be circular in shape and the pore size (diam-
eter) was determined using the circle area formula.
3D view of cylindrical structures was provided by a
stereoscopic microscope Leica M205 C (Leica Microsys-
tems GmbH).
Porosity was determined individually for every pros-
thesis. Two prosthesis ends, 1cm in length were cut off
and weighted on analytical lab scale. Afterward, the sam-
ples (n = 2 for each prosthesis) were glued upright to
SEM stubs and coated with 15nm of gold as described
above. SEM images (n = 6) at magnification 200 × were
taken, and n = 30 wall thickness measurements were
made for each sample to determine individual layers’
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
thickness and total wall thickness. e results (sample
weight (ms) and total wall thickness (
δs)
)were averaged
and used to determine prostheses porosity (
ε)
using for-
mula:
ε
=1−
m
s
δs•Ss
ρp
∗
100%
, where
ρp
is a density of
polyurethane ChronoFlex®C75A,
ρp
=
1.2g
/
cm3
[28],
Ss
is a sample’s side surface determined using for-
mula
Ss
=
2π(r
+
δ)L
, where r is a prosthesis inner
Fig. 1 A Layers arrangement in Nano and Micro prosthesis, (B) Cross section of Nano and Micro prosthesis’ wall, (C) stereoscopic image of Nano
and Micro prosthesis, (D) macroscopic image of prostheses (Nano and Micro mix)
Table 1 SBS process parameters applied for each layer in Nano and Micro prostheses. The layer that is present in a given prosthesis
type is marked with “ + ”, a layer that is absent is marked with “‑ “
Layer Nano Micro Polymer
conc. [%w/w]Solution
vol. [ml] Collector-nozzle tip
distance [cm] Collector
rotational speed
[rpm]
Solution ow
rate [ml/h] Air ow
rate [MPa]
nanofibers + ‑ 2 3 50 3 000 30 0.1
dense microfibers + + 5 6 10 3 000 30 0.1
microfibers + + 5 20 30 3 000 30 0.1
aligned microfibers + + 5 4 30 20 000 30 0.1
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
radius, r = 0.25cm and L is a sample length L = 1cm. e
results are presented as mean value ± SD.
Mechanical properties
Prostheses of 5cm in length (n = 5 for each type) were
placed in the pneumatic jaws of the testing machine
Instron 3345 equipped with 50kN static load cell. Pros-
theses were stretched at the rate of 10mm/min until the
break. Dedicated Bluehill software automatically deter-
mined maximum load, elongation at break, Young’s mod-
ulus, and ultimate tensile stress. e results are presented
as a mean value ± SD.
Leakage anddelamination tests
e leakage test was carried out as follows: prostheses
of 4cm length (n = 3 for each type) were mounted in
a closed flow system connected to a peristaltic pump
Zalipm PP1B-05A (Zalipm) and 0.9% NaCl solution was
circulated in the system (through the prosthesis) for
1h at a flow rate of 20ml/min. During the test, sam-
ples were checked for any signs of leakage through the
prostheses’ walls. After the leakage test, prostheses were
dried at 20°C for 24 h. en, the samples were glued
to SEM stubs with conductive carbon adhesive tape and
covered with 15nm layer of gold. Materials cross-sec-
tions were analyzed using scanning electron microscopy
Phenom G1.
e above-described flow system was also used to test
the permeability of the prostheses’ walls in contact with
blood. Freshly drawn whole blood was connected to the
flow system and the prostheses were perfused for 1h at
a flow rate of 20ml/min. During that time, macroscopic
observations were carried out to assess whether there is
any blood leakage through the prosthesis’ wall.
Additionally, a static delamination test was carried out.
Prostheses of 1.5cm length (n = 3 for each type of pros-
theses and for each timepoint) were prepared and placed
in 1.5 ml Eppendorf® test tubes fully filled with 0.9%
NaCl solution. Test tubes were closed and placed in an
incubator at 37°C for 7, 14, or 30days. After this time,
the prostheses were dried at 20°C for 24h and investi-
gated using scanning electron microscopy Phenom G1.
Hemocompatibility ofmaterials
Blood tests were performed using fresh human blood
from healthy volunteers. Blood was collected in 1.8ml
test tubes containing citrate (BD Vacutainer, Franklin
Lakes, NJ, USA).
Static platelet adhesion
For static analysis, round shape samples (n = 2 for each
type of material) were placed in 24-well plate with the
luminal surface of the prosthesis facing up. In order to
stabilize and flatten the material, each sample was placed
in CellCrown (Sigma-Aldrich) inserts. Subsequently,
500µl of 0.9% NaCl solution in ultrapure water was added
to wells with samples and plate was incubated at 37°C for
30min. en, NaCl solution was removed and 200µl of
platelet-rich plasma (PRP) was added to every well con-
taining the samples. PRP was prepared using two “slow”
centrifugations: 150 g for 14 min (first centrifugation)
and 150g for 12min (second centrifugation). e plate-
let density in PRP was 1 × 106 platelets/µL. Plate with
materials was incubated at 37°C for 90min. Next, PRP
was removed, and samples were thoroughly rinsed with
0.9% NaCl to remove blood residues. Finally, samples
were prepared for SEM analysis. Briefly, materials were
incubated in 4% paraformaldehyde for 24h at 4°C. Next,
the samples were dehydrated by 10min immersion steps
in 50, 60, 70, 80, 90, and 100% ethanol solution (EtOH),
and for 20 min in 1:2 hexamethyldisilazane:ethanol
(HMDS:EtOH), 2:1 HDMS:EtOH and 100% HDMS solu-
tion. Finally, the samples were glued to SEM stubs with
conductive carbon adhesive tape (luminal surface of
prostheses up) and covered with 15nm layer of gold. e
% of platelet-coated area was counted from SEM images
of every sample, taken at 3000 × magnification. Addition-
ally, pictures at magn. = 5000 × were taken in order to
present the morphology of surface-adhered platelets in
detail. e platelet adhesion assay was done in triplicate,
with change of blood donor each time. For every sample
n = 10 SEM images were taken. e average values for all
materials were calculated from 60 images (10 images × 3
experiments × 2 samples).
In this assay, PTFE was cut from vascular prosthesis
(FlowLine Bipore, Jotec) and used as a reference material
that induces low platelet adherence.
Hemolysis
Round samples with diameter of 14mm (n = 3 for each
type of prosthesis) were placed in 48-well plate with the
luminal surface of the prosthesis facing up. To separate
erythrocytes from plasma, fresh blood was centrifuged
at 700g for 5min and plasma was removed from blood
tubes. en, erythrocytes were diluted 20 × in ultracold
PBS and 500µl of erythrocyte suspension was added to
wells with materials. PBS was used as a negative control
and 0.2% TritonX-100 as a positive control. Triplicate
samples were placed on a shaker at 300rpm for 1h, at
37°C. Afterward, 600µl of solution from every well was
centrifuged at 700g for 1min, and 200µl of supernatant
was transferred triplicate to 96-well plate. e absorb-
ance at 540nm was measured using a plate reader Epoch
Biotek (Biokom).
Hemolysis rate was calculated using the following
formula:
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
where: AS – sample absorbance, ACP – mean positive con-
trol absorbance, C– mean negative control absorbance.
Results are presented as mean hemolysis rate ± SD.
Endothelial cell culture
Human umbilical vein endothelial cells (HUVECs) were
isolated from freshly collected umbilical cords (kindly
provided by the Dept. of Gynaecology, University Hos-
pital Erlangen) and grown in supplemented endothelial
cell growth medium (EGM-2, Promo Cell, Germany).
Accutase solution was used for cell harvesting. Cells from
passages 1 or 2 were used in experiments. All experi-
ments were repeated 3 times, in each experiment the
material was used in duplicate. e use of human mate-
rial was approved by the local ethics committee at the
University Hospital Erlangen (review number 14-85_3-B
from 01.02.2022).
Static cell seeding onat materials – 2D model
Flat samples were cut off from cylindrical grafts, steri-
lized with 70% ethanol, washed with sterile PBS, and
placed in 24 well cell culture inserts. en, materials were
seeded with HUVECs (5 × 104 cells/sample) and incu-
bated at 37°C for 1, 3, and 7days. Culture media were
changed 24h after seeding and then every second day.
To analyze cell viability, Alamar Blue assay was per-
formed according to manufacturer’s protocol. Briefly,
after 1, 3, or 7days of cell culture, materials with cells
growing on the surface were transferred to a new 24-well
plate and gently washed with sterile PBS. en Alamar
Blue working solution was added to each well (500 µl/
well) and incubated with samples at 37°C for 18h in the
dark. e fluorescence of the Alamar Blue solution was
measured at Ex./Em = 550/590 nm using a plate reader
(SpectraMax iD3, Molecular Devices).
Magnetic cell seeding oncylindrical prostheses – 3D model
Cell seeding was also performed on cylindrical vascu-
lar prostheses. For this, all materials were cut to equal
length of 5cm. Samples were sterilized with 70% etha-
nol, washed with sterile PBS, and placed in transparent
cell culture tubes. 1% agarose solution was used to fix the
prostheses in a vertical position inside the cell culture
tubes. Before cell seeding prostheses were preincubated
with EGM-2 medium for at least 1h.
HUVECs were seeded on the lumen of the pros-
theses using magnetic seeding technique. Cells were
pre-incubated with superparamagnetic iron oxide
nanoparticles (SPIONs) in cell culture flasks for 24 h
at 37 °C as described before [29]. After incubation,
HR
=
A
S−
A
CN
A
CP
−A
CN
∗
100%
the SPION-loaded cells were harvested and counted.
HUVECs were suspended in the culture media and trans-
ferred into the luminal space of each prosthesis (1 × 106
cells/prosthesis). Immediately after transferring the cell
suspension, the scaffolds were exposed to a radially sym-
metric magnetic field for 15 min using the VascuZell
endothelizer (Vascuzell Technologia S.L., Madrid, Spain).
e cell culture tubes with prostheses were then carefully
removed from the endothelizer and placed in the incuba-
tor for 1, 3, or 7days. e culture medium was changed
24h after seeding and then every second day.
Cell staining andimage analysis
After the given cell culture period, cells growing on the
lumen surface were fixed with 4% buffered paraformal-
dehyde (Roth GmbH, Karlsruhe, Germany) and permea-
bilized with 0.2% Triton X-100 (Sigma-Aldrich, Munich,
Germany) in PBS. F-actin filaments were stained by
Alexa488-phalloidin (Invitrogen, ermo Fisher) and
visualized using fluorescence microscope Zeiss Axio
Observer Z1 (Zeiss, Jena, Germany) at 10 × magnifica-
tion. To observe cells growing inside cylindrical prosthe-
ses, the materials were cut along the longitudinal axis,
pressed to the glass slides, and then visualized using mul-
tiple mode (2 × 5). Cell counting was performed using
the ImageJ software (Fiji, version 1.47v).
Data analysis andstatistical analysis
2D cell culture model andplatelet adhesion assay
Cell coverage was calculated in 6 circular samples with a
diameter of 11mm (2 replicates × 3 independent experi-
ments). For each sample, at least 3 microscopic images
(magnification = 10x) were taken in randomly selected
places and the cell coverage was calculated for every
image. e average coverage was then calculated for each
sample and the resulting boxplot was based on these
6 average values for all 6 samples. A boxplot in a%-b%
range indicates that in a given group of materials, there
was at least one sample with a% coverage and at least one
sample with b% coverage.
3D cell culture model
Cell coverage was calculated in 5 cylindrical samples
(diameter 6mm, length 5cm) from 3 independent exper-
iments). For each sample, at least 2 multi-tile scan micro-
scopic images (magnification = 10x) were taken. Each
“tile” represents the standard analysis area at 10 × and the
multi-tile scans covered the surface of 2 tiles (prosthesis
circumference) × 5 tiles (prosthesis length), correspond-
ing to an area of approx. 1.3mm × 4.5mm. For each sam-
ple, 2 multi-tile scans were performed and the results
were averaged. Based on 5 averaged values for all 5 sam-
ples a boxplot was plotted in a%-b% range, indicating
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
that in a given group of materials, there was at least one
sample with a% coverage and at least one sample with b%
coverage.
e results of the other measurements (mechanical
analysis, delamination assay, hemolysis) were presented
as mean values ± SD. Statistical significance ofdifferences
was analyzed using single-factor ortwo-factor analysis of
variance (ANOVA) for p < 0.05 with post-hoc Tukey’s test
(OriginPRO 2020b).
Results
Morphology ofprostheses
In the first part of this study, we produced and compared
two types of vascular prostheses. e wall of prostheses
consisted of several layers, each of which should fulfill
specific functions. A schematic diagram comparing the
arrangement of layers in the respective types of prosthe-
ses is shown in Fig.1A, while Fig.1B shows SEM pictures
of their wall cross-sections. Essentially, the two scaffold
types differed by the presence of nanofibrous layer on
the luminal side of the Nano prosthesis. Our earlier stud-
ies have shown that prostheses made only of microfib-
ers are leaky. erefore, we decided to include alayer of
densely arranged microfibers in both types of prostheses.
is layer, whose thickness was about 10% of the total
wall thickness acted as a sealing. e thickness of this
layer was selected as a result of our previous work (data
not shown), and the prostheses’ permeability was tested
in a flow system using saline (see below). e next layer
to the outside is a layer of loosely arranged microfibers,
intended mainly to ensure appropriate mechanical prop-
erties of the prostheses (e.g., flexibility) and to achieve
the desired wall thickness. is is the thickest layer of the
graft, constituting about 80% of the total wall thickness.
e thin outermost layer, representing about 10% of the
total wall thickness consists of circumferentially aligned
fibers and is designed to promote attachment of SMCs.
Figure 1C and Fig.1D show microscopic (stereoscopic
microscopy) and macroscopic photos of prostheses. e
macroscopic appearance of both types of prostheses was
similar.
e evaluation of fiber diameter on internal surfaces of
both types of prostheses is presented in Fig. 2. Average
fiber diameters of luminal surfaces of Nano and Micro
prostheses were 262 ± 68nm and 991 ± 251nm, respe c-
tively. Additionally, pore size measurements were per-
formed on the luminal surface of prostheses. Average size
of pores was 2.5 ± 0.9µm for Nano and 3.7 ± 1.7 µm for
Micro scaffold. It must be noted that in Micro prosthe-
sis, the fibrous areas covered only about 14% of luminal
surface.
e described SBS process allowed to obtain prosthe-
ses with comparable properties, but with differing lumen
topography. e luminal surface of Nano prostheses was
characterized by nanofibers with single defects (beaded
fibers) present on the surface (as indicated by the arrows
in the Fig.2B), while the luminal surface of Micro pros-
theses had a more heterogenous structure, includ-
ing areas of smooth solid surface with pores and small
fibrous areas with large, flattened fibers.
Mechanical properties ofprostheses
As shown in Fig. 3A, the mechanical properties of
Nano and Micro prostheses were similar. No significant
Fig. 2 A Fiber diameter distribution and (B) internal surface morphology (arrows indicate defects present on the surface) for internal surface of
Nano and Micro prostheses
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
Fig. 3 AMechanical properties of Nano and Micro prosthesis (n = 5), (B) Load‑extension curve for Nano and Micro prostheses. “*” indicates a
change‑point related to the rapture of the two outer microfiber layers, # indicates a change‑point related to the rapture of the internal dense
microfiber layer
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
differences regarding wall thickness were observed
between the two types of prostheses. e total wall thick-
ness was 698 ± 44 µm for Nano and 680 ± 45 µm for
Micro scaffolds. e thickness of innermost nanolayer
for Nano prostheses was 9 ± 2µm, whereas the thickness
of dense microfibers’ layer was 88 ± 9µm for Nano and
97 ± 12µm for Micro.
In accordance with this, no significant differences were
detected in mechanical properties of the prostheses. Both
types presented elastic behavior with high elongation at
break values. Porosity was similar and equaled 43 ± 10%
for Nano and 40 ± 8% for Micro. Young’s modulus values
for Nano and Micro prostheses were 2.5 ± 0.2 MPa and
2.4 ± 0.1 MPa, respectively, while the respective maxi-
mum load values were 58.3 ± 3.8 N and 61.2 ± 3.2 N. Ulti-
mate tensile stress for porous sample was 10.9 ± 1.8 MPa
for Nano and 10.0 ± 0.8MPa for Microprostheses. Elon-
gation at break value was lower for Nano (407 ± 46%)
than for Micro prostheses (478 ± 30%), but the difference
was not statistically significant. Figure3B shows a typi-
cal load-extension curve. e shape of the curve is simi-
lar for both types of prostheses. e first change-point
(marked as “*”) in the curve is related to the rupture of
the two outer microfibrous (aligned and non-aligned)
layers of the prosthesis. e test ended when the remain-
ing layer (dense microfibers) was ripped up (the second
change-point marked as “#”).
Prosthesis leakage anddelamination test
During 1h contact between Nano or Micro prostheses
and 0.9% NaCl solution in flow system, no soaking or
leakage was observed. No leakage was also observed dur-
ing blood contact analysis in flow system. Additionally,
no delamination of layers was observed either after 1 h
contact with 0.9% NaCl solution in a flow system during
dynamic delamination test, or after 30days of incubation
in 0.9% NaCl (static delamination test). Representative
SEM images of Nano and Micro prostheses, showing
their cross-sections after 30 days of static incubation
in 0.9% NaCl solution, are presented in Fig.4A. Cross-
section SEM images of Nano and Micro prostheses after
dynamic(1h) and static(7 and 14 days) analysis are pre-
sented in supplementary data. Figure 4B presents the
results of wall thickness measurements before and after
7, 14, and30days of static delamination. No significant
changes in wall thickness were observed, regardless of
duration of the test and the type of prosthesis.
Biological evaluation
e detailed characterization of produced scaffolds dem-
onstrated that Nano and Micro prostheses differ only in
their luminal surface morphology. In the second part of
this study, we, therefore, evaluated the influence of this
structural difference on hemocompatibility and endothe-
lial cell attachment to the produced scaffolds.
Platelet adhesion
e luminal surface of the materials after the plate-
let adhesion test is shown in Fig.5A. e percentage of
the platelet-occupied area is shown in Fig.5B. e aver-
age values obtained for all tested materials were similar
and no statistically significant differences were detected
(p > 0.05 for all pairs). However, SEM images pointed to
the differences in the morphology of the adherent plate-
lets. On the Nano surfaces, platelet aggregates formed
strongly flattened structures. ere was a relatively large
variation in platelet coverage between samples, ranging
from 1 to 19% and the average value of platelet coverage
was 8.6%. In the case of Micro materials, a different mor-
phology of the adhered platelets was observed. e cells
formed relatively large aggregates, which had a spheri-
cal, rounded form. Highly flattened aggregates were rare.
e variation in platelet coverage values between the
samples was similar to Nano, in the range of 2–17% and
the average value of platelet coverage was 6.2%. e aver-
age platelet coverage values for both types of prosthe-
ses were close to those observed on the surface of PTFE
(7.0%). In the case of PTFE, the adherent platelets formed
highly flattened layer and no spherical aggregates were
observed.
Hemolysis
e results presented in Table 2 demonstrated that
hemolysis rate upon blood contact with Nano or Micro
prostheses was < 1%. e produced prostheses thus do
not cause blood hemolysis.
Endothelial cell culture
Static seeding onat materials – 2D model
e results of cell culture on flat samples (the 2D model)
are shown in Fig.6. Microscopic analysis showed that the
cells showed the correct morphology and adhered to the
surface of the fibers (Fig.6A). Starting from the first day
of culture (D1) a higher percentage of cell-covered area
(Fig.6B) was detected on Nano surfaces, but the differ-
ences were not statistically significant. Cellular coverage
for the Nano surface was in the range of 5–30%, with an
average value of 20%. For Micro surfaces, cellular cover-
age values ranged from 5 to 20%, with an average of 13%.
A similar relationship was obtained on the third day of
culture (D3). In the case of Nano surfaces, the cellular
coverage values were higher and ranged from 25 to 65%,
with an average value of 42%. For the Micro surface, the
values ranged from 20 to 55%, with an average value of
35%. On the 7th day of culture (D7), the differences
between the cellular coverage values for Nano and Micro
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
surfaces increased and were in the range of 60–85% and
20–65%, respectively. e average endothelial cell cover-
age was 73% for Nano and was significantly larger than
for Micro samples (43%, p < 0.05).
Cell viability analysis (Fig. 6C) confirmed the micro-
scopic observations. At each time point, the fluorescence
value was higher for Nano than Micro surfaces.
Magnetic seeding oncylindrical material – 3D model
e results of cell seeding in the 3D cylindrical scaffolds
are shown in Fig.7. Both, morphology of the ECs (Fig.7A)
and the cell coverage values (Fig. 7B) indicate that no
significant differences were observed between Nano
and Micro prostheses. It is worth emphasizing that cell
growth was highly heterogeneous, especially in the later
days of culture. On the 7th day of culture, both types of
prostheses showed areas of cells forming monolayer-like
spots, but there were also areas without any adherent
cells. Generally, in both cases, the cellular coverage values
were significantly lower than in the 2D culture, being in
the range of 5%—30% for Nano and 5%- 25% for Micro
prostheses. ere were also large differences in the values
obtained between multiplicate experimental samples.
Discussion
Previous studies on small-diameter vascular grafts have
revealed that layered structure of vascular prostheses
significantly improves their mechanical properties and
better mimics the structure and functions of native
blood vessel [30]. Each layer of such prosthesis should
Fig. 4 A Prostheses cross‑section SEM images after 30 days of static delamination tests, (B) Prostheses wall thickness before and after static
delamination test
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
fulfill certain requirements to enable vessel multifunc-
tionality including anti-thrombogenic function, inhibi-
tion of intimal hyperplasia, reduction ofinflammation,
and enhancement endothelialization after implantation.
Many research groups have previously developed
multilayered prostheses, however, most commonly each
layer was produced by a different method and often
from different polymers. For instance, Yuan etal. cre-
ated prostheses in which inner, middle,and outer layers
were produced by ink printing, wet spinning, andelec-
trospinning, respectively. e authors claimed that only
the combined use of the 3 methods allowed for the pro-
duction of prostheses with the desired wall thickness
and mechanical properties [31]. By combining E-jet
Fig. 5 AMorphology of surface‑adhered platelets (magn. = 5 kx) and (B) percentages (n = 6) of platelet‑occupied area for Nano and Micro surfaces.
PTFE was used as a low‑thrombogenic reference material
Table 2 Hemolysis rate of Nano and Micro prostheses
Nano Micro
Hemolysis rate [% of positive control] 0.4 ± 0.1 0.1 ± 0.1
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
3D printing and electrospinning methods, Huang etal.
produced tri-layered prostheses, which exhibited bet-
ter mechanical properties in comparison to electrospun
monolayer grafts [30]. Generally, fibrous constructs are
very popular in vascular engineering, due to their 3D
structure properties mimicking extracellular matrix.
Moreover, they offer the possibility to customize fibrous
scaffold surface properties (fiber diameter, fiber align-
ment, material porosity) during production, depending
on the requirements of the selected cell type [32].
In this study, multi-step solution blow spinning of med-
ical grade PU enabled us to produce two types of layered
fibrous vascular prostheses that differ in their luminal
surface morphology. Both types of prostheses were made
of three main layers. e outer layer, identical to the
Micro and Nano type, was made of microfibers with an
average diameter of 1µm. is layer was designed to sup-
port the development of SMCs that build the walls of
native blood vessels. In our previous studies, the growth
of SMCs on fibers with average diameters in the range
of 200, 500, and 900µm was analyzed [27], showing that
SMC growth on fibrous scaffolds with fiber diameters
of ~ 1µm is improved in comparison to smaller diame-
ters. In addition, other studies suggested that not only the
size but also the orientation of fibers supports the pro-
cess of SMC attachment and growth [33, 34]. Based on
those results, the outer surfaces of both types of prosthe-
ses were designed to contain homogeneous, circumfer-
entially oriented microfibers with average fiber diameter
of ~ 1µm.
A similar diameter was selected to produce a middle
layer of the prostheses, composed of loose, non-aligned
Fig. 6 Cell culture with flat materials: (A) HUVECs growth, (B) cell coverage and (C) cell viability after 1,3 and 7 days of culture (n = 6)
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
microfibers. e main task of this ~ 560µm thick inter-
mediate layer was to ensure the appropriate mechanical
properties, which are an important factor determining
the success of grafts upon implantation. Invivo, blood
vessels are constantly exposed to pulsatile pressure and
undergo constant deformation [35], so the prostheses
should be produced from highly elastic materials. Fur-
thermore, differences in mechanical properties between
the implanted prosthesis and the native vessel can lead to
aneurysm formation [36] or anastomotic intimal hyper-
plasia [37]. Mechanical properties of small-diameter vas-
cular grafts should therefore be similar to properties of
native vessels, which they are intended to replace (e.g.
coronary artery), or to the commonly used autografts,
such as saphenous vein, with longitudinal elastic modulus
about 24MPa and ultimate tensile stress about 6MPa, or
internal thoracic (mammary) artery, with longitudinal
elastic modulus about 17MPa and ultimate stress about
4MPa [38]. In this study, maximum load values of about
60 N and ultimate tensile stress values of about 10MPa
confirmed high mechanical strength of the produced
scaffolds. e prostheses had Young’s modulus values
of about 2.5MPa, which proves their elasticity. It must
be noted that Young’s modulus of electrospun polyure-
thane prostheses strongly depends on the type of poly-
mer used, e.g. values reported for Cardiomat were below
1MPa [39] and for Tecothane around 6MPa [40]. Grasl
etal. reported electrospun Pellethane prostheses with an
average fiber diameter of about 900nm and axial Young’s
modulus reaching 10MPa [41].
e mechanical properties of prostheses change with
the change in the average diameter of the fibers that build
their walls [42]. As the morphology and thickness of the
intermediate layer were the same for Nano and Micro
prostheses, it was expected that their mechanical prop-
erties will be comparable. e additional nanofiber layer
in the Nano-type prostheses was only about 10µm thick
and had therefore no significant effect on the mechanical
properties of the entire prosthesis.
Generally, porous structures that mimic extracellu-
lar matrix provide a suitable microenvironment for cell
growth and tissue regeneration. However, in the case of
vascular grafts, they pose a risk of leakage [43]. To over-
come this problem, a low-porosity, impermeable com-
pact layer made of densely arranged microfibers was
added during fabrication by changing the nozzle-collec-
tor working distance, so that the resulting final porosity
of the wall was about 40%. is approach allowed us to
effectively prevent the leakage as demonstrated in the
closed flow system perfusion tests.
e main goal of this study was to evaluate whether the
change in the morphology of the internal surface of the
prosthesis, without the change in its mechanical properties,
Fig. 7 Cell culture with cylindrical materials: (A) HUVECs growth and
(B) cell coverage after 1, 3, and 7 days of culture (n = 5)
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Łopianiaketal. Journal of Biological Engineering (2023) 17:20
has a significant impact on the adhesion of platelets and
ECs. Previous studies indicated that cell growth on unmodi-
fied polymers including PU is at most moderate and that
chemical modifications of the surface, e.g. by introducing
peptides, are necessary in order to create a stable layer of
the endothelium [44, 45]. On the other hand, many studies
reported a strong influence of fiber diameter on cell adhe-
sion. Taking this into account, we analyzed the adhesion of
platelets and ECs on the internal surfaces of the prostheses.
e adhesion of blood platelets to the inner surface of the
prosthesis is an undesirable phenomenon that may lead
to the formation of a clot and thrombotic occlusion of the
prosthesis lumen. It is also known that the process of plate-
let adhesion can be influenced by the physical properties of
the surface, i.e. roughness [46], topography [47], sub-micron
texturing [48]. Studies of platelet adhesion to a solvent-
cast film coated with electrospun nanofibers made from
poly[acrylonitrile-co-(N-vinyl-2-pyrrolidone)] (PANCNVP)
[49] demonstrated that while the platelets did not adhere
to the surface of the film, they did adhere to the surface of
the nanofibers. In the micro range, however, Milleret etal.
reported that electrospun PU scaffolds with smaller fiber
diameters (< 1 μm) reduce platelet adhesion [50]. Authors
stated that not only size of the surface features (e.g. fibers),
but also differences in roughness are very likely responsi-
ble for the differential platelet adhesion. Our study showed
that changing the morphology of the internal surface of
the prostheses within the range reported here had a negli-
gible effect on the total percentage of platelet-covered sur-
face. Interestingly, however, the change in luminal surface
morphology did influence the morphology of the adherent
platelets. Nano-type surface promoted the strong flattening
of platelets and their aggregates, while on Micro-type sur-
faces mostly spherical clusters were formed. Such a differ-
ence in the morphology of the adherent platelets may affect
the level of their activation and, as a result, the probability
of thrombosis. In the context of thrombogenicity, it is worth
emphasizing that platelet adhesion to our PU prostheses
was overall comparable to the PTFE, which is considered
a low-thrombogenic material. e ChronoFlex PU used in
this study is also characterized by low platelet adhesiveness
and has been successfully used in the production of artificial
heart, among others.
One of the key aspects of successful small-diameter
vascular grafting is a rapid endothelialization of prosthe-
ses [7]. Endothelial monolayer lining the inner surface of
arteries, veins and capillaries constitutes a barrier between
blood and tissues [51]. Furthermore, vascular endothelium
controls and regulates blood flow. Also, an intact and tight
endothelium prevents platelet activation, adhesion, and
aggregation. is helps to maintain the patency of vascu-
lar graft after implantation [52]. In this study, we hypoth-
esized that changing the prostheses’ luminal surface
morphology by introducing layer of nanofibers would
enhance the EC attachment and ability to form mon-
olayer. Similar effect was previously reported by Chung
et al. who increased roughness of the smooth PU films
by grafting PU chains with different molecular weights
and chain lengths, showing that increased nanoscale sur-
face roughness enhances the adhesion and growth of ECs
[53]. Furthermore, studies of endothelial cell growth on
the surface of PLC/collagen fibers with diameters of 0.27,
1, 2.39, 4.45µm showed that cells grown on 0.27µm fib-
ers formed strong focal adhesion, whereas cells grown on
2.39 and 4.45µm fibers presented a spindle-shaped mor-
phology with very few focal adhesion points [42]. In our
study, the process of cell colonization in 2D (flat samples,
cell seeding by sedimentation) was faster on Nano-type
surfaces. e difference in the percentage of cell-covered
area between the Micro and Nano prostheses was particu-
larly evident in the later days of the culture. After 7days
of culture, the cellular coverage on the Nano surfaces was
in the range of 60–85% and the cell growth was relatively
uniform on the entire analyzed surface. On Micro sur-
faces, the cell coverage values were not only lower, but
the cell growth was also patchy and non-uniform. is is
certainly related to the greater heterogeneity of the surface
morphology of the Micro type, which likely translates into
non-uniform cell growth. To investigate whether nanofi-
brous surface morphology promotes endothelialization
of 3D constructs, we employed the magnetic cell seeding
method, which was successfully used to populate other
types of cylindrical biomaterials [29, 54]. However, the
obtained cell coverage values were overall lower than in
2D samples (30% after 7days of culture) and did not sig-
nificantly differ between both Nano and Micro prostheses.
is trend was observed for both types of materials, for all
observation time points. e mechanisms of this effect are
unclear thus far but may be related to the differences in the
seeding process between 2 and 3D samples. In the case of
the 2D model, the cell suspension was applied to the upper
surface of the flat samples placed in CellCrown inserts and
incubated for 24h, allowing the cells to sediment onto the
material under the influence of gravity. In the 3D model,
cylindrical samples placed in a vertical position in cell cul-
ture tubes were filled with cell suspension and exposed to
magnetic field for 15min. Consequently, the period when
the cells had a chance to adhere to the luminal, cylindri-
cal surface of the prosthesis was much shorter in case of
3D samples that were subsequently placed in the incuba-
tor and remained in a vertical position during the whole
culture time. us, the cells that did not effectively attach
to the luminal surface during the 15min exposure to the
magnetic field would have fallen to the bottom of the cul-
ture tube. Expectedly, gravity did not work in favor of cell
adhesion in vertically placed scaffolds leading to relatively
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Page 15 of 17
Łopianiaketal. Journal of Biological Engineering (2023) 17:20
poor attachment of endothelial cells to the cylindrical wall
of 3D model and explaining the differences in endothelial
coverage between 2 and 3D samples.
In summary, multilayered cylindrical prostheses pro-
duced from medical-grade PU by solution blow spinning
method, represent a promising alternative to autologous
vessels or synthetic polymers. While differing in luminal
surface morphology, the designed prostheses showed a
high elasticity, good mechanical strength, and a platelet
adhesion level comparable to PTFE. Changing the lumi-
nal surface morphology by adding a nanofibrous layer
significantly improved endothelialization of the flat sam-
ples. However, this morphological enhancement was
not strong enough to show a significant effect during
colonization of the entire cylindrical prostheses, which
is a more demanding process. us, in order to achieve
successful cell colonization of 3D cylindrical prosthe-
ses, it will be necessary to introduce additional chemical
modifications of their surface (such as introduction of
bioactive endothelial cell-selective adhesive molecules,
e.g. REDV, IKAV) to overcome current limitations and
improve endothelialization efficacy.
Conclusions
is study aimed to develop non-thrombogenic small-
diameter vascular grafts with appropriate mechanical
properties and to evaluate the influence of luminal sur-
face morphology on scaffold hemocompatibility and
endothelial cell attachment. Collectively, our data dem-
onstrate that multistep solution blow spinning method
allows to produce cylindrical structures with layers of tai-
lorable thickness and porosity, whose mechanical prop-
erties conform to small-diameter vascular grafts. e
developed prostheses did not cause hemolysis in contact
with blood and there was no significant difference in the
percentage of platelet-covered area for Nano and Micro
surfaces. Nanofibrous surfaces promoted stronger adhe-
sion of platelets and their aggregates, resulting in the
presence of flattened structures. On the contrary, Micro
surfaces were characterized by the presence of spherical
aggregates, which indicates their weaker adhesion. is
variation in surface-adhered platelets might indicate dif-
ferences in their activation level.
Endothelial coverage after 1, 3, and 7days of 2D cul-
ture was higher on Nano prostheses. However, this effect
was not seen in 3D culture, where cylindrical prostheses
were colonized using magnetic seeding method. Taken
together, the produced scaffolds meet the material and
mechanical requirements for vascular prostheses, but
their biological properties must be further improved to
enhance endothelialization efficiency.
Abbreviations
ECs Endothelial cells
EtOH Ethanol
HMDS Hexamethyldisilazane
HUVECs Human umbilical vein endothelial cells
PBS Phosphate buffer saline
PET Polyethylene terephthalate
PRP Platelet‑rich plasma
PTFE Polytetrafluoroethylene
PUs Polyurethanes
SBS Solution blow spinning
SEM Scanning electron microscopy
SMCs Smooth muscle cells
Supplementary Information
The online version contains supplementary material available at https:// doi.
org/ 10. 1186/ s13036‑ 023‑ 00337‑9.
Additional le1:Figure. Cross‑sectional SEM images of Nano and Micro
prostheses after static (7 and 14 days) and dynamic (1h) delamination test.
Authors’ contributions
I.Ł.: methodology, investigation, data analysis, writing—original draft, visualiza‑
tion, W. Rz.: investigation, M. C.: investigation, data analysis, T.C.: supervision,
project administration, funding acquisition, writing—review & editing, I.C.:
methodology, supervision, project administration, funding acquisition, writ‑
ing—review & editing, B.B‑R.: conceptualization, methodology, investiga‑
tion, data analysis, writing—original draft, visualization, supervision, project
administration, funding acquisition. All authors read and approved the final
manuscript.
Funding
The work was supported by the National Science Centre, Poland (grant no.
UMO‑2020/39/I/ST5/01131), the Deutsche Forschungsgemeinschaft (DFG,
German Research Foundation; grant no. CI 162/4–1), the National Centre
for Research and Development, Poland (grant no. LIDER/18/0104/L‑8/16/
NCBR/2017), Faculty of Chemical and Process Engineering, Warsaw University
of Technology (grant agreement no. 504/04627) and the Polish National
Agency for Academic Exchange (NAWA) under The Bekker Programme, (grant
no. BPN/BEK/2021/1/00434/U/00001).
Availability of data and materials
The datasets generated during the study are available from the corresponding
author upon reasonable request.
Declarations
Ethics and approval and consent to participate
The use of human material was approved by the local ethics committee at the
University Hospital Erlangen (review number 14‑85_3‑B from 01.02.2022).
Consent for publication
Not applicable.
Competing interests
The authors declare no competing interests.
Author details
1 Faculty of Chemical and Process Engineering, Warsaw University of Technol‑
ogy, Waryńskiego 1, 00‑645 Warsaw, Poland. 2 Doctoral School of Warsaw
University of Technology, Warsaw University of Technology, Pl. Politechniki 1,
00‑661 Warsaw, Poland. 3 Section of Experimental Oncology Und Nanomedi‑
cine (SEON), Else Kröner‑Fresenius‑Stiftung‑Professorship, ENT‑Department,
Universitätsklinikum, Erlangen, Germany. 4 Centre for Advanced Materials
and Technologies CEZAMAT, Warsaw University of Technology, Poleczki 19,
02‑822 Warsaw, Poland.
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Page 16 of 17
Łopianiaketal. Journal of Biological Engineering (2023) 17:20
Received: 3 January 2023 Accepted: 5 March 2023
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