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Received: 28 March 2025
Revised: 24 April 2025
Accepted: 27 April 2025
Published: 2 May 2025
Citation: Mohammed, A.;
Tirnoveanu, A.; Webb, W.R.; Melaibari,
A.A.; Memi´c, A.; Aslam, M.; Elshaer,
A.; Hassanin, H.; Essa, K. In Vitro
Characterization of 3D-Printed
PLA/CPO Oxygen Releasing
Scaffolds: Mechanical and Biological
Properties for Bone Tissue
Engineering. J. Manuf. Mater. Process.
2025,9, 149. https://doi.org/
10.3390/jmmp9050149
Copyright: © 2025 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/).
Article
In Vitro Characterization of 3D-Printed PLA/CPO Oxygen
Releasing Scaffolds: Mechanical and Biological Properties for
Bone Tissue Engineering
Abdullah Mohammed 1,2 , Alice Tirnoveanu 3, William Richard Webb 3, Ammar A. Melaibari 2,4 ,
Adnan Memi´c 2, Mohammad Aslam 5, Amr Elshaer 6, Hany Hassanin 7, * and Khamis Essa 1, *
1School of Engineering, University of Birmingham, Birmingham B15 2TT, UK; axm1777@student.bham.ac.uk
2Centre of Nanotechnology, King Abdulaziz University, Jeddah P.O. Box 80204, Saudi Arabia;
aamelaibari@kau.edu.sa (A.A.M.); amemic@kau.edu.sa (A.M.)
3Institute of Medical Sciences, Faculty of Medicine, Health and Social Care, Canterbury Christ Church
University, Canterbury CT1 1QU, UK; alice.tirnoveanu@canterbury.ac.uk (A.T.);
w.r.webb1@icloud.com (W.R.W.)
4Mechanical Engineering Department, Faculty of Engineering, King Abdulaziz University (KAU),
Jeddah P.O. Box 80204, Saudi Arabia
5Centre of Excellence in Environmental Studies, King Abdulaziz University,
Jeddah P.O. Box 80204, Saudi Arabia; magmuhammad21@kau.edu.sa
6Drug Discovery, Delivery and Patient Care (DDDPC), School of Life Sciences, Pharmacy and Chemistry,
Kingston University London, Kingston Upon Thames KT1 2EE, UK; a.elshaer@kingston.ac.uk
7School of Engineering, Canterbury Christ Church University, Canterbury CT1 1QU, UK
*Correspondence: hany.hassanin@canterbury.ac.uk (H.H.); k.e.a.essa@bham.ac.uk (K.E.)
Abstract: The addition of oxygen-releasing biomaterials into 3D-printed scaffolds presents
a novel approach to enhancing bone scaffolds, yet no
in vitro
studies have demonstrated
the effect of oxygen-generating filaments on scaffold biological and mechanical properties.
This study introduces a polylactic acid (PLA)/calcium peroxide (CPO) composite fila-
ment, designed for oxygen release, which is a key factor for early-stage bone regeneration.
The PLA/CPO composite filament was fabricated via wet-mixing, solvent evaporation,
and hot-melt extrusion, followed by fused deposition modeling (FDM) with optimized
parameters to achieve high structural fidelity (25% porosity, 0.60mm pore size).
In vitro
characterization, including mechanical, morphological, and biological assessments, demon-
strated that, post-cell culturing, mechanical strength improved, which indicates improved
scaffold resilience. The scaffold exhibited gradual oxygen release over a 3-day period, and
gene expression analysis confirmed notable upregulation of osteogenic markers RUNX2,
SPP1, and SP7 in vitamin D-supplemented conditions. The mechanical strength improved
from approximately 2.8 MPa in the control group to 5.0 MPa in scaffolds cultured with
osteogenic media. This study provides the first
in vitro
evidence that oxygen-releasing
3D-printed filaments can improve both mechanical properties and biological response in
scaffolds, demonstrating the functional integration of sustained oxygen delivery, enhanced
mechanical properties, and increased osteogenic activity in a single 3D-printed scaffold.
Keywords: 3D printing; oxygen-generating scaffolds; PLA/CPO composite; bone tissue
regeneration; fused deposition modeling (FDM)
1. Introduction
The rising need for organ transplants and tissues has led to a severe shortage of
available donors, creating a significant gap between supply and demand. This shortage
J. Manuf. Mater. Process. 2025,9, 149 https://doi.org/10.3390/jmmp9050149
J. Manuf. Mater. Process. 2025,9, 149 2 of 14
represents a major biomedical challenge, with over 106,000 individuals in the United States
currently waiting for a transplant. According to data from March 2022, this shortage has
tragic consequences, as 17 people die each day while waiting for a transplant due to delays
in receiving the necessary organs [
1
]. Bones are among the most highly sought-after tissues
for transplantation in the United States, and finding a donor match is a significant challenge
for bone graft procedures [
2
]. Autografts, which utilize tissue from the patient’s own
body, are widely considered the best option for bone fracture repair and regeneration. This
method promotes healing by using the patient’s natural tissue, minimizing the risk of
rejection. However, autografts are not always practical, particularly in cases where the
bone defect is too large to repair with the available tissue or when highly precise shaping is
required, as in complex facial bone surgeries. In such situations, alternative solutions must
be explored to achieve successful outcomes [
3
]. Tissue engineering has become a viable
alternative to bone grafting for bone regeneration. This method involves the creation of a
bone scaffold that includes growth factors, stem cells, and biocompatible and biodegradable
materials, which can aid in bone fracture healing and enhance the incorporation of the
graft [
4
,
5
]. The scaffolds are specifically designed to provide essential structural support
while promoting tissue regeneration. By mimicking the natural extracellular matrix, they
create an environment that encourages cell growth and repair, ultimately assisting the
tissue in recovering its functional capabilities [6–8].
Recently, the application of 3D printing technologies for producing bone scaffolds has
attracted significant interest. This approach allows for the creation of scaffolds with specific
external designs and porous internal structures, enabling the development of scaffolds with
customized functionality [
9
,
10
]. Among the various 3D printing methods, fused deposition
modeling (FDM) is frequently employed in tissue engineering due to its affordability,
accessibility, and ease of use [
11
]. Moreover, FDM achieves a printing accuracy of up to
+/
−
0.5 mm and can utilize a wide range of biocompatible and biodegradable polymeric
materials suitable for tissue engineering [11,12].
Tissue engineering makes use of a variety of materials to repair bone, tendon, and
skin. Among these, polycaprolactone (PCL), polylactic acid (PLA), and polyglycolide or
poly-glycolic acid (PGA) are particularly useful due to their favorable mechanical and
biochemical properties [
13
,
14
]. PLA is a material recognized for its beneficial physical
and mechanical properties, alongside its biocompatibility and biodegradability. These
attributes are significantly influenced by factors such as its molecular weight. Higher
molecular weights generally enhance strength and durability [
15
]. Together, these factors
determine how well PLA performs in various applications, especially in medical and
tissue engineering contexts [
16
]. These properties make PLA an excellent candidate for a
wide variety of industrial applications, particularly in the development of medical devices.
Additionally, PLA is compatible with FDM 3D printers and has been approved by the US
Food and Drug Administration (FDA) for several biomedical uses [
17
]. However, while
tissue engineering has shown promising results in laboratory settings, its clinical application
has been largely limited to treating small tissue defects, typically measuring just a few
millimetres. This limitation stems from the difficulty in achieving adequate vascularization,
which is crucial for supplying the necessary oxygen to support tissue survival and growth.
Without sufficient blood vessel formation, larger tissue constructs struggle to receive the
oxygen and nutrients they need, hindering their effectiveness in clinical settings [
18
,
19
]. The
insufficient oxygen levels in engineered tissues are a significant barrier to bone regeneration
and the overall effectiveness of scaffolds. This shortage restricts the growth of cells that
attach to the scaffold, which is essential for successful tissue development [20].
Researchers have designed scaffolds that release oxygen by utilizing various solid
peroxide particles, such as magnesium peroxide, calcium peroxide, and sodium percar-
J. Manuf. Mater. Process. 2025,9, 149 3 of 14
bonate. The usual process involves the disintegration of these particles in water, resulting
in the release of oxygen through hydrolysis, as displayed in Equations (1) and (2) [
7
,
21
].
Calcium peroxide (CPO) is frequently selected as an oxygen-releasing agent because it is
both economical and widely available in the market, making it a practical choice compared
to other materials [22–24].
CaO2+2H2O→Ca(OH)2+H2O2(1)
2H2O2→O2+ 2H2O (2)
Studies suggest that the fabrication of bone scaffolds incorporating 3D-printed oxy-
genation filaments can significantly enhance bone tissue regeneration and healing. The
incorporation of an oxygen source within the scaffold has been found to promote vascular-
ization and enhance the scaffold’s efficacy [
19
]. The aim of this study is to investigate the
in vitro
performance of a newly developed oxygen-releasing PLA/CPO scaffold fabricated
using FDM. This work is the first to assess the combined effects of sustained oxygen release,
post-culture mechanical behavior, and osteogenic gene expression in bone stem cells.
2. Materials and Methods
2.1. Materials
PLA filament with a diameter of 1.75 mm was purchased from the Shenzhen eSUN
Industrial Co., Ltd. (Shenzhen, China). For the oxygen-releasing component, calcium
peroxide (CPO) was obtained from Sigma-Aldrich (St. Louis, MO, USA); it has a particle size
of 200 mesh and a purity of 75%. Additionally, catalase, derived from bovine serum with an
activity level of 5000 units/mg, was also purchased from Sigma-Aldrich. Dichloromethane
(DCM), a solvent used in various applications, was acquired from the same supplier, along
with deionized water for use in the experiments.
2.2. PLA/CPO Filament
PLA filaments (20 g) were first cut into small pieces and dissolved in 100 mL of DCM
at room temperature for 30 min, using a magnetic stirrer set at 700 rpm. Once the PLA was
fully dissolved, CPO powder was added to the PLA and the mixture was vigorously stirred
for an additional 90 min to ensure thorough incorporation. The resulting homogeneous
solution was then poured into a large plate and allowed to dry for 24 h. After drying, the
composites were cut into smaller pieces suitable for loading into the hot melt extruder. A
custom single-screw extruder with a nozzle diameter of 2 mm was employed to extrude
the composite materials at a nozzle temperature of 140
◦
C and an extrusion speed of
2.5 cm/s. Figure 1shows the schematic diagram of the process and the extruded filament
under the microscope. The prepared PLA-CPO filaments were immersed in a phosphate-
buffered saline (PBS) solution (pH 7.4), which mimics the pH and osmolarity of human
fluids. Additionally, 10 mg of catalase was also added into the solution to catalyze the
decomposition of hydrogen peroxide (H
2
O
2
) into water and oxygen, ensuring accurate
detection of the oxygen-releasing capability of the material. Oxygen levels were measured
in triplicate using an oxygen sensor to track the oxygen release on daily use.
J. Manuf. Mater. Process. 2025,9, 149 4 of 14
J. Manuf. Mater. Process. 2025, 9, x FOR PEER REVIEW 4 of 14
Figure 1. (a) Schematic diagram of the PLA-CPO composite lament process (b) the extruded la-
ment under SEM.
2.3. 3D Printing of Bone Scaolds
A commercial Fused Deposition Modeling (FDM) 3D printer, specically the Creality
Ender 3 Pro, manufactured by the Shenzhen Creality 3D Technology Co., Ltd. in China,
was utilized to fabricate the bone scaolds under a range of parameters outlined in Table
1. The scaolds were designed in a square shape, measuring 8 × 8 mm in both length and
width, with a height of 1.5 mm. Each scaold was engineered to have a porosity of 25%,
featuring pores with a size of 0.60 mm. This design was intended to optimize the scaolds
for tissue regeneration by providing adequate support and permeability for cell growth.
Table 1. 3D printer parameters used to print PLA scaolds.
Sample No.
Printing Temperature
(°C)
Building Platform
Temperature (°C)
Printing Speed
(mm/s)
1
200
70
25
2
200
70
50
3
200
70
75
2.4. Characterizations
2.4.1. Morphological Analysis
The scaolds were designed in a square shape, measuring 8 × 8 mm in both length
and width, with a height of 1.5 mm. Each scaold was engineered to have a porosity of
25%, featuring pores with a size of 0.60 mm. This design was intended to optimize the
scaolds for tissue regeneration by providing adequate support and permeability for cell
growth. The micrograph images were used for morphological analysis using ImageJ
1.53f51, an open-source image analysis software developed by the National Institutes of
Health (NIH) in the USA. The contact angle was measured on printed scaold samples
Figure 1. (a) Schematic diagram of the PLA-CPO composite filament process (b) the extruded filament
under SEM.
2.3. 3D Printing of Bone Scaffolds
A commercial Fused Deposition Modeling (FDM) 3D printer, specifically the Creality
Ender 3 Pro, manufactured by the Shenzhen Creality 3D Technology Co., Ltd. in China,
was utilized to fabricate the bone scaffolds under a range of parameters outlined in Table 1.
The scaffolds were designed in a square shape, measuring 8
×
8 mm in both length and
width, with a height of 1.5 mm. Each scaffold was engineered to have a porosity of 25%,
featuring pores with a size of 0.60 mm. This design was intended to optimize the scaffolds
for tissue regeneration by providing adequate support and permeability for cell growth.
Table 1. 3D printer parameters used to print PLA scaffolds.
Sample No. Printing
Temperature (◦C)
Building Platform
Temperature (◦C)
Printing Speed
(mm/s)
1 200 70 25
2 200 70 50
3 200 70 75
2.4. Characterizations
2.4.1. Morphological Analysis
The scaffolds were designed in a square shape, measuring 8
×
8 mm in both length
and width, with a height of 1.5 mm. Each scaffold was engineered to have a porosity of
25%, featuring pores with a size of 0.60 mm. This design was intended to optimize the
scaffolds for tissue regeneration by providing adequate support and permeability for cell
growth. The micrograph images were used for morphological analysis using ImageJ 1.53f51,
an open-source image analysis software developed by the National Institutes of Health
J. Manuf. Mater. Process. 2025,9, 149 5 of 14
(NIH) in the USA. The contact angle was measured on printed scaffold samples using a
KRÜSS DSA30E drop shape analyzer (Germany). The sessile drop method was employed
with 10
µ
L deionized water droplets, dispensed at a rate of 2
µ
L/s. Measurements were
performed at room temperature and averaged over three replicates persample. The scaffold
morphology was examined using a JEOL JSM-6010LA scanning electron microscope. Prior
to imaging, samples were sputter-coated with a 5 nm gold layer using a Quorum Q150R ES
sputter coater to improve surface conductivity. Digital images of the drops were analyzed
by computer software (DSA4, version 2.1). The water droplets were placed on relatively
flat scaffold regions. Given the porous nature of the scaffold, minor liquid bridging across
interconnected pores was possible; however, multiple measurements were performed across
different areas of the scaffold to mitigate this effect. The final contact angle values were
averaged to ensure repeatability and accuracy. The data analysis focused on comparing
contact angle variations at different printing speeds, correlating surface topology changes
with material deposition behavior during printing.
2.4.2. Osteogenic Differentiation In Vitro
Cell Culture:
Human mesenchymal stem cells (Lonza) were cultured in Dulbecco’s Modified Ea-
gle’s Medium—GlutaMax, a nutrient-rich medium designed to support cell growth. This
medium was supplemented with 5% (v/v) human platelet-rich plasma (PRP), which pro-
vides essential growth factors, as well as 1% (v/v) non-essential amino acids to promote
cellular health and function. Additionally, 1% (v/v) penicillin-streptomycin-amphotericin
was included to prevent bacterial and fungal contamination. The cells were maintained in
this enriched environment for 2 weeks, allowing them to expand and prepare for subse-
quent seeding in experimental applications.
Post-expansion, the PLA rods were sterilized for 24 hrs in 100% ethanol and 1 hr
post-sterilization the PLA rods were pretreated by soaking in PBS containing 10ng/mL
fibronectin at 37
◦
C. After 1 hour, the rods were washed twice with sterile PBS, and then
4
×
10
5
cells were seeded onto each rod. The rods were then placed in a ST180 PLUS
CO2 incubator (Benchmark Scientific, Darmstadt, Germany) for 20 min to facilitate cell
attachment. After this incubation, the rods were submerged in KOSR media for 24 h. This
media was formulated with Dulbecco’s Modified Eagle’s Medium—GlutaMax, enhanced
by 10% (v/v) KnockOut™ Serum Replacement, along with 1% (v/v) non-essential amino
acids and 1% (v/v) penicillin-streptomycin-amphotericin, ensuring an optimal environment
for the cells to flourish. After 24 hrs, the rods were separated into two groups: Control
group (KOSR without vitamin D), and experimental differentiation group (where the media
was supplemented with vitamin D, Beta-glysophosphate, and ascorbic-2-phosphate).
In the negative control plate, the medium did not contain vitamin D. The media were
changed twice weekly for both the control and differentiation group over 21 days. PLA-only
scaffolds were not included as a control group. The experimental aim was to evaluate
osteogenic responses of the PLA/CPO scaffold under different biological stimulation
conditions. Each group was tested in triplicate (n= 3), and the results were presented
descriptively without statistical testing, reflecting the exploratory nature of the study.
RNA Extraction:
The rods were then placed in a ST180 PLUS CO
2
incubator (Benchmark Scientific,
Darmstadt, Germany) for 20 min to facilitate cell attachment. After this incubation, the
rods were submerged in KOSR media for 24 h. This media was formulated with Dul-
becco’s Modified Eagle’s Medium—GlutaMax, enhanced by 10% (v/v) KnockOut™ Serum
J. Manuf. Mater. Process. 2025,9, 149 6 of 14
Replacement, along with 1% (v/v) non-essential amino acids and 1% (v/v) penicillin-
streptomycin-amphotericin, ensuring an optimal environment for the cells to flourish.
RNA Quantification:
The extracted RNA was quantified by absorbance reading at a 260/280 nm ration with
Tecan i-200-Pro. The measurements were performed in triplicate. The mean concentration
value per sample was used for further calculation in the qRT-PCR set-up.
Quantitative Real-Time Reverse Transcriptase Polymerase Chain Reaction (qRT-PCR)
For the analysis of gene expression related to Glyceraldehyde 3-phosphate dehydroge-
nase (GAPDH), Runt-related transcription factor 2 (RUNX2), and Collagen type I alpha
1 chain (COL1A1), Secreted Phosphoprotein 1 (SPP1), and Sp7 transcription factor (SP7),
the Qiagen QuantiNova SYBR Green RT-PCR kit was used, according to the supplier’s
guidelines.
Customized primer sets for GAPDH, RUNX2, COL1A1, SPP1 and SP7 were previously
designed and verified against the human cell line MG63 (Osteosarcoma) supplied by ATCC.
All primer sequences were designed using human gene data from the Ensembl genome
browser and the NCBI Gene database. The specificity of the alignments was evaluated
using the NCBI Primer-BLAST tool. The RT-PCR reaction was organized as follows:
The fold change value was further analyzed against the GAPDH expression and
vitamin D was normalized utilizing delta-delta Ct to calculate the fold change against the
control sample KOSR.
2.4.3. Mechanical and Microstructural Properties
To evaluate the mechanical properties of the filaments, a universal Instron 3367 testing
machine (Norwood, MA, USA) equipped with a 30 kN load cell was employed. The
testing involved tensile testing of the filament with diameters ranging from 1.75 mm to
1.95 mm and a length of 90 mm. The filaments were securely held in place using manual
grips, and the machine’s crosshead speed was maintained at a constant rate of 5 mm/min.
Each experiment was conducted three times, and the average values were calculated for
accuracy. The results are presented as the mean
±
standard deviation (SD). An Ultima
IV X-ray diffractometer (XRD) (Rigaku, Tokyo, Japan) equipped with Cu K
α
radiation
and referenced with the ICDD (PDF-2/release 2011 RDB) database, including DB card
No. 01-071-4107, was used to analyze the extruded filaments. The measurements were
conducted at a goniometer speed of 1.00 sec per step with a step size of 0.100◦.
2.5. Statistical Analysis
Each experiment was conducted in triplicate to ensure reliability, and the mean value
was calculated from these trials. The results are presented as averages with the
±
standard
deviation (SD) to show variability. Data analysis and visualization were performed using
Origin software (OriginPro 8.0, Origin Lab Inc., Northampton, MA, USA), which helped
create clear graphical representations of the findings.
3. Results and Discussion
3.1. 3D Printing of Scaffolds
During extrusion, filaments containing 6% CPO exhibited optimal extrudability, form-
ing smooth and uniform filaments suitable for 3D printing. In contrast, higher CPO
concentrations led to increased brittleness, compromising both the printability and struc-
tural integrity. Therefore, a CPO concentration of 6% was maintained across all experiments
to ensure consistent material performance and reproducibility in the biological evaluations.
The release of oxygen of the PLA/CPO filament is shown in Figure 2. The oxygen release
J. Manuf. Mater. Process. 2025,9, 149 7 of 14
profile demonstrates a gradual increase in oxygen concentration over 3 days. It shows
that the embedded CPO undergoes hydrolysis at a steady rate. While the oxygen release
results cover a short period (days), bone tissue regeneration is a long process that takes
weeks to months. However, the evidence of early oxygenation provided by the PLA/CPO
can play a key role in angiogenesis and osteogenic differentiation, which are fundamental
steps in bone healing. The early release ensures that cells survive and proliferate in the
initial hypoxic environment, supporting early vascularization before nutrition and the
blood supply establishes. Further studies are needed to optimize oxygen release duration
to extend beyond 3 days and sustain oxygen delivery throughout the healing process [
25
].
J. Manuf. Mater. Process. 2025, 9, x FOR PEER REVIEW 7 of 14
oxygen release results cover a short period (days), bone tissue regeneration is a long pro-
cess that takes weeks to months. However, the evidence of early oxygenation provided by
the PLA/CPO can play a key role in angiogenesis and osteogenic dierentiation, which
are fundamental steps in bone healing. The early release ensures that cells survive and
proliferate in the initial hypoxic environment, supporting early vascularization before nu-
trition and the blood supply establishes. Further studies are needed to optimize oxygen
release duration to extend beyond 3 days and sustain oxygen delivery throughout the
healing process [25].
Figure 2. (a) Image of oxygen released during the degradation process shown as bubbles from the
laments (b) Oxygen release of the PLA/CPO laments over a 3-day period.
Figure 3 presents optical images of the 3D-printed PLA/CPO scaolds with both a
side view and a top view of the scaold structure. The gure demonstrates the pores and
surface features of the printed scaold sample. The distinct layers also highlight how the
printing parameters, such as speed, contribute to the scaold’s vertical integrity. At lower
speeds, scaolds typically show uneven layer heights, but at higher speeds, as represented
in Figure 3a, the layers appear smoother and more uniform, correlating with the ndings
on accuracy. The 3D printer parameters were mainly determined by two variables: tem-
perature and speed of printing. While maintaining the temperature constant at ~200 °C,
the printing speed was set to three dierent levels (Table 1). Figure 3b shows the pores
are regularly spaced and exhibit a high degree of uniformity. This uniformity was more
pronounced at higher printing speeds, such as 75 mm/s, where the accuracy reached
98.33%. The well-formed pores, critical for nutrient exchange and cell inltration in tissue
engineering, align with the structural requirements for bone scaolds. At lower speeds,
the pores tend to show overlap or deformation. The optical images serve as visual proof
of the scaold’s dimensional accuracy and porosity. The pore interconnectivity and size
suggest that the PLA/CPO composite scaold is likely to facilitate eective cell prolifera-
tion and dierentiation, matching the requirements outlined in the literature.
Figure 3. Optical images of the 3D printed scaolds (a) side image and (b) top image with lighting
source showing porosity.
Figure 2. (a) Image of oxygen released during the degradation process shown as bubbles from the
filaments (b) Oxygen release of the PLA/CPO filaments over a 3-day period.
Figure 3presents optical images of the 3D-printed PLA/CPO scaffolds with both
a side view and a top view of the scaffold structure. The figure demonstrates the pores
and surface features of the printed scaffold sample. The distinct layers also highlight
how the printing parameters, such as speed, contribute to the scaffold’s vertical integrity.
At lower speeds, scaffolds typically show uneven layer heights, but at higher speeds, as
represented in Figure 3a, the layers appear smoother and more uniform, correlating with
the findings on accuracy. The 3D printer parameters were mainly determined by two
variables: temperature and speed of printing. While maintaining the temperature constant
at ~200
◦
C, the printing speed was set to three different levels (Table 1). Figure 3b shows
the pores are regularly spaced and exhibit a high degree of uniformity. This uniformity was
more pronounced at higher printing speeds, such as 75 mm/s, where the accuracy reached
98.33%. The well-formed pores, critical for nutrient exchange and cell infiltration in tissue
engineering, align with the structural requirements for bone scaffolds. At lower speeds, the
pores tend to show overlap or deformation. The optical images serve as visual proof of the
scaffold’s dimensional accuracy and porosity. The pore interconnectivity and size suggest
that the PLA/CPO composite scaffold is likely to facilitate effective cell proliferation and
differentiation, matching the requirements outlined in the literature.
Moreover, Mota et al. noted that scaffolds with regular pore distribution support
enhanced tissue integration, which can be inferred from the top view in Figure 3. The
consistency of the pore shapes and the absence of irregularities further corroborate the
improvements in accuracy and quality at higher printing speeds, reinforcing the conclusions
drawn from previous sections of the analysis [26].
J. Manuf. Mater. Process. 2025,9, 149 8 of 14
J. Manuf. Mater. Process. 2025, 9, x FOR PEER REVIEW 7 of 14
oxygen release results cover a short period (days), bone tissue regeneration is a long pro-
cess that takes weeks to months. However, the evidence of early oxygenation provided by
the PLA/CPO can play a key role in angiogenesis and osteogenic differentiation, which
are fundamental steps in bone healing. The early release ensures that cells survive and
proliferate in the initial hypoxic environment, supporting early vascularization before nu-
trition and the blood supply establishes. Further studies are needed to optimize oxygen
release duration to extend beyond 3 days and sustain oxygen delivery throughout the
healing process [25].
Figure 2. (a) Image of oxygen released during the degradation process shown as bubbles from the
filaments (b) Oxygen release of the PLA/CPO filaments over a 3-day period.
Figure 3 presents optical images of the 3D-printed PLA/CPO scaffolds with both a
side view and a top view of the scaffold structure. The figure demonstrates the pores and
surface features of the printed scaffold sample. The distinct layers also highlight how the
printing parameters, such as speed, contribute to the scaffold’s vertical integrity. At lower
speeds, scaffolds typically show uneven layer heights, but at higher speeds, as represented
in Figure 3a, the layers appear smoother and more uniform, correlating with the findings
on accuracy. The 3D printer parameters were mainly determined by two variables: tem-
perature and speed of printing. While maintaining the temperature constant at ~200 °C,
the printing speed was set to three different levels (Table 1). Figure 3b shows the pores
are regularly spaced and exhibit a high degree of uniformity. This uniformity was more
pronounced at higher printing speeds, such as 75 mm/s, where the accuracy reached
98.33%. The well-formed pores, critical for nutrient exchange and cell infiltration in tissue
engineering, align with the structural requirements for bone scaffolds. At lower speeds,
the pores tend to show overlap or deformation. The optical images serve as visual proof
of the scaffold’s dimensional accuracy and porosity. The pore interconnectivity and size
suggest that the PLA/CPO composite scaffold is likely to facilitate effective cell prolifera-
tion and differentiation, matching the requirements outlined in the literature.
Figure 3. Optical images of the 3D printed scaffolds (a) side image and (b) top image with lighting
source showing porosity.
The 3D printing of bone scaffolds using PLA/CPO composite material (Figure 3)
was investigated to achieve the highest accuracy and quality. The 3D printer parameters
were mainly determined by two variables: temperature and speed of printing. While
maintaining the temperature constant at ~200
◦
C, the printing speed was set to three
different levels (Table 1). The dimensional accuracy of the printed scaffolds at different
speeds was determined based on the number of pores fabricated without overlapping or
defacing the original design.
Figure 4presents a detailed analysis of the accuracy, quality, and surface characteristics
of the 3D-printed PLA/CPO scaffolds at various printing speeds (25 mm/s, 50 mm/s, and
75 mm/s). This figure provides quantitative data and visual insights into how the printing
speed influences key aspects of scaffold formation, including the dimensional accuracy,
slope angles, layer line width, and rough surface. The graph in Figure 4a shows a linear
relationship between accuracy and printing speed. The graph in Figure 4a demonstrates
a clear linear relationship between the printing speed and dimensional accuracy of the
scaffolds. As the speed increases from 25 mm/s to 75 mm/s, the accuracy improves
significantly, with the highest accuracy of 98.33% achieved at 75 mm/s. At 25 mm/s,
the accuracy is drastically lower, around 48%, which indicates that at lower speeds, the
printed pores are either deformed or overlap due to excessive material deposition or nozzle
movement inconsistencies. At 50 mm/s, the accuracy improves but remains suboptimal
compared to the highest speed. The optimal accuracy at 75 mm/s indicates that a faster
printing speed allows for more precise material deposition without compromising the pore
structure.
Figure 4b highlights how the printing speed affects the layer line width of the printed
scaffolds. At lower speeds of 25 mm/s, the layer line width is inconsistent and irregular,
contributing to a rougher surface. However, at 75 mm/s, the width becomes much more
uniform, with a measured width of 0.40 mm. The slope angle, which indicates the angle
between the layers, is much steeper at lower speeds (about 18
◦
at 25 mm/s). This steeper
angle leads to less even material distribution. As the speed increases, the slope angle
decreases to 11.5
◦
at 75 mm/s, resulting in a smoother and more stable surface structure,
crucial for scaffold integrity. Typically, smoother layer transitions and reduced slope angles
at higher speeds improve scaffold functionality by enhancing structural homogeneity.
Figure 4c compares the water contact angles of scaffolds printed at different speeds,
which reflects surface morphology. The contact angle increases with speed, showing 96.5
◦
at 25 mm/s, 97
◦
at 50 mm/s, and 103
◦
at 75 mm/s. A higher contact angle indicates a
smoother and more hydrophobic surface. At higher speeds, the smoother surface reduces
the roughness, thus increasing the contact angle. This suggests improved surface quality,
as rougher surfaces typically lead to lower contact.
J. Manuf. Mater. Process. 2025,9, 149 9 of 14
J. Manuf. Mater. Process. 2025, 9, x FOR PEER REVIEW 9 of 14
some irregularities remain, suggesting moderate improvements in layer deposition and
surface roughness. At the highest speed, Figure 4f, the SEM image displays a much
smoother, more even surface. The layers are well aligned, with minimal defects or ridges.
This significant improvement in surface morphology at high speed confirms the trend ob-
served in both the accuracy and water contact angle measurements. Generally, a smoother
surface will have a higher contact angle, while a rougher surface will have a lower contact
angle [27]. Furthermore, it is clear from Figure 4d–f that the surface roughness becomes
smoother and more symmetrical as the printing speed is increased. It was reported by
Wangwang while printing an object with 100% infill that printing at high speed can cause
cavities and delamination at the interface between different layers [28], which is opposite
to our findings. It can be argued that our case is different from what has been reported in
the literature because our composite material has a higher viscosity compared to pure
PLA, making it more viscous due to the embedded CPO particles within the PLA matrix
[29]. Hence, printing parameters for high accuracy and quality can vary depending on the
materials used, as they are highly influenced by the materials’ properties. As a result with
the PLA/CPO composite, a higher extruding speed is required to achieve the desired ac-
curacy and quality.
Figure 4. Evaluation of 3D printing accuracy and quality, (a) accuracy percentages of printed scaf-
folds at different printing speeds, (b) slope degree and width of scaffolds layer line at different
printing speeds, and (c) water contact angles of printed scaffolds surface at different printing speeds.
SEM images of printed scaffolds layer at different printing speeds of: (d) 25 mm/s, (e) 50 mm/s, and
(f) 75 mm/s, (n = 3).
3.2. Gene Expression
Skeleton can be either formed by intramembranous ossification or endochondral os-
sification. Mesenchymal cells differentiate directly into osteoblasts, which lead to the for-
mation of intramembranous bones [30]. On the other hand, the replacement of the carti-
laginous structure by bone will form endochondral bones. Around 90% of bone is made
of type 1 collagen [31]. The laer is a triple helix structure made of 2 type 1 collagen strand
(Col1a1) and on Col1a2 strand [32]. Furthermore, Col1a1 and Col1a2, as well as various
transcription factors are involved in bone development and maintenance. One important
transcription factor is RUNX2, which is essential for osteoblast differentiation and bone
Figure 4. Evaluation of 3D printing accuracy and quality, (a) accuracy percentages of printed scaffolds
at different printing speeds, (b) slope degree and width of scaffolds layer line at different printing
speeds, and (c) water contact angles of printed scaffolds surface at different printing speeds. SEM
images of printed scaffolds layer at different printing speeds of: (d) 25 mm/s, (e) 50 mm/s, and
(f) 75 mm/s, (n= 3).
The SEM image in Figure 4d shows significant surface roughness and irregular layer
formation. The material appears uneven, with visible ridges and cavities. This highlights
the negative effect of low-speed printing, where excess material accumulates, creating a
rough, uneven surface. While the surface becomes smoother than at 25 mm/s, Figure 4e,
some irregularities remain, suggesting moderate improvements in layer deposition and sur-
face roughness. At the highest speed, Figure 4f, the SEM image displays a much smoother,
more even surface. The layers are well aligned, with minimal defects or ridges. This signifi-
cant improvement in surface morphology at high speed confirms the trend observed in both
the accuracy and water contact angle measurements. Generally, a smoother surface will
have a higher contact angle, while a rougher surface will have a lower contact angle [
27
].
Furthermore, it is clear from Figure 4d–f that the surface roughness becomes smoother
and more symmetrical as the printing speed is increased. It was reported by Wangwang
while printing an object with 100% infill that printing at high speed can cause cavities
and delamination at the interface between different layers [
28
], which is opposite to our
findings. It can be argued that our case is different from what has been reported in the
literature because our composite material has a higher viscosity compared to pure PLA,
making it more viscous due to the embedded CPO particles within the PLA matrix [
29
].
Hence, printing parameters for high accuracy and quality can vary depending on the
materials used, as they are highly influenced by the materials’ properties. As a result
with the PLA/CPO composite, a higher extruding speed is required to achieve the desired
accuracy and quality.
3.2. Gene Expression
Skeleton can be either formed by intramembranous ossification or endochondral ossifi-
cation. Mesenchymal cells differentiate directly into osteoblasts, which lead to the formation
of intramembranous bones [
30
]. On the other hand, the replacement of the cartilaginous
structure by bone will form endochondral bones. Around 90% of bone is made of type 1
collagen [
31
]. The latter is a triple helix structure made of 2 type 1 collagen strand (Col1a1)
and on Col1a2 strand [
32
]. Furthermore, Col1a1 and Col1a2, as well as various transcription
J. Manuf. Mater. Process. 2025,9, 149 10 of 14
factors are involved in bone development and maintenance. One important transcription
factor is RUNX2, which is essential for osteoblast differentiation and bone formation [
33
].
Another significant factor is phosphoprotein 1 (Spp1), also known as osteopenia, produced
by osteoblasts. Spp1 plays a crucial role in regulating bone mineralization and remodeling
processes [
34
]. SP7, another transcription factor, is involved in osteoblast differentiation
and mineralization [
35
]. In the present experimental configuration, the combination of
scaffold composition and culture conditions led to an enhancement in the gene expression
of Human Mesenchymal stem cells (hMSCs), cells associated with bone development in
contrast to those grown in a monolayer. The upregulation of RUNX2 reflects the early-
stage commitment of hMSCs to the osteoblast lineage. SP7, acting downstream of RUNX2,
supports osteoblast maturation, while SPP1 (osteopontin) is a late marker involved in
extracellular matrix organization and mineralization. The observed co-activation of these
markers suggests progression through multiple stages of osteogenic differentiation.
In Figure 5, the gene expression results were normalized to the control filaments which
were cultured in the absence of vitamin D. When Human Mesenchymal stem cells (hMSCs)
were differentiated in the presence of vitamin D upregulation of the bone transcription
factors essential for bone differentiation, (RUNX2, SPP1, and SP7) were observed. Con-
versely, a downregulation of COL1A1 was noted (Figure 5). Notably, a distinct increase
was detected in the expression of RUNX2 genes in hMSC cells cultured, in contrast to
the expression levels of COL1a1, SPP1, and SP7 genes, as well as the literature [
36
]. The
downregulation of COL1A1 under vitamin D stimulation may reflect a biological shift
from matrix production to mineralization, a phenomenon reported during later osteogenic
stages. These results indicate that the oxygen-releasing scaffold, in combination with vita-
min D, fosters an environment conducive to both early (RUNX2) and late (SPP1) osteogenic
differentiation phases.
J. Manuf. Mater. Process. 2025, 9, x FOR PEER REVIEW 10 of 14
formation [33]. Another significant factor is phosphoprotein 1 (Spp1), also known as oste-
openia, produced by osteoblasts. Spp1 plays a crucial role in regulating bone mineraliza-
tion and remodeling processes [34]. SP7, another transcription factor, is involved in oste-
oblast differentiation and mineralization [35]. In the present experimental configuration,
the combination of scaffold composition and culture conditions led to an enhancement in
the gene expression of Human Mesenchymal stem cells (hMSCs), cells associated with
bone development in contrast to those grown in a monolayer. The upregulation of RUNX2
reflects the early-stage commitment of hMSCs to the osteoblast lineage. SP7, acting down-
stream of RUNX2, supports osteoblast maturation, while SPP1 (osteopontin) is a late
marker involved in extracellular matrix organization and mineralization. The observed
co-activation of these markers suggests progression through multiple stages of osteogenic
differentiation.
In Figure 5, the gene expression results were normalized to the control filaments
which were cultured in the absence of vitamin D. When Human Mesenchymal stem cells
(hMSCs) were differentiated in the presence of vitamin D upregulation of the bone tran-
scription factors essential for bone differentiation, (RUNX2, SPP1, and SP7) were ob-
served. Conversely, a downregulation of COL1A1 was noted (Figure 5). Notably, a dis-
tinct increase was detected in the expression of RUNX2 genes in hMSC cells cultured, in
contrast to the expression levels of COL1a1, SPP1, and SP7 genes, as well as the literature
[36]. The downregulation of COL1A1 under vitamin D stimulation may reflect a biological
shift from matrix production to mineralization, a phenomenon reported during later os-
teogenic stages. These results indicate that the oxygen-releasing scaffold, in combination
with vitamin D, fosters an environment conducive to both early (RUNX2) and late (SPP1)
osteogenic differentiation phases.
Figure 5. Delta-Delta Ct values for hMSCs differentiated in the presence of vitamin D (n = 3).
3.3. Mechanical Properties
The XRD in Figure 6a shows the microstructural in the PLA/CPO composite after
extrusion. Following hot extrusion, the PLA peak is wide and significantly less intense,
indicating a less/semi-crystalline phase. The CPO peaks are more visible, confirming that
CPO is still present. CPO particles are observed as bright areas dispersed within the PLA
0.01
0.10
1.00
10.00
100.00
1000.00
RUNX 2 COL 1A1 SPP 1 SP7
Fold Change (Log10)
Gene
Relative qRT-PCR Vitamin D normalised to KOSR control
Figure 5. Delta-Delta Ct values for hMSCs differentiated in the presence of vitamin D (n= 3).
3.3. Mechanical Properties
The XRD in Figure 6a shows the microstructural in the PLA/CPO composite after
extrusion. Following hot extrusion, the PLA peak is wide and significantly less intense,
indicating a less/semi-crystalline phase. The CPO peaks are more visible, confirming that
CPO is still present. CPO particles are observed as bright areas dispersed within the PLA
J. Manuf. Mater. Process. 2025,9, 149 11 of 14
matrix; see Figure 6b. The image shows relatively small CPO particles with a more uniform
distribution, though some CPO particles tend to agglomerate in certain areas.
J. Manuf. Mater. Process. 2025, 9, x FOR PEER REVIEW 11 of 14
matrix; see Figure 6b. The image shows relatively small CPO particles with a more uni-
form distribution, though some CPO particles tend to agglomerate in certain areas.
Figure 6. (a) X-ray diffraction (XRD) of PLA/CPO before and after extrusion, (b) Stress strain dia-
gram of filaments containing no cells (gray), with KOSR (blue) and with vitamin D, (c) SEM image
shows white areas corresponding to CPO particles and grey areas representing the PLA matrix.
The stress–strain diagram in Figure 6b provides a comparison of the mechanical be-
havior of PLA/CPO composite filaments under different conditions: No Cells (black line),
KOSR (red line), and D (vitamin D-blue line). The figure shows how osteogenic differen-
tiation influences filaments’ mechanical properties. The No Cells control filament, which
was left in solution for 2 weeks without cell integration, exhibits the lowest mechanical
performance, with a tensile strength of approximately 2.8 MPa and a strain of ~0.7%. The
early failure and limited ductility indicate that PLA/CPO degradation weakened the fila-
ment’s structure. On the other hand, PLA/CPO cultured with KOSR and D differentiation
media exhibit a significant increase in both tensile strength and strain, confirming the re-
inforcing effect of cell integration. The KOSR filaments demonstrate the highest mechan-
ical improvement, with a tensile strength of ~5 MPa and a strain of ~1.6%, outperforming
both the D and the No Cells samples. The almost twofold increase in strength and strain
for both cell-cultured filaments highlights the role of osteogenic differentiation in improv-
ing the samples’ flexibility and structural resilience. These improvements are likely at-
tributed to early extracellular matrix (ECM) deposition by differentiating cells. As cells
adhere to the scaffold and begin to remodel their microenvironment, secreted collagen
and other matrix proteins form a cohesive interface that enhances the mechanical integ-
rity. Such biological reinforcement is characteristic of early-stage osteogenesis and reflects
the scaffold’s potential to support functional tissue development.
A key observation is that the KOSR filaments maintain strength over a larger strain
range before ultimate failure. This behavior indicates that osteogenic cells within the
KOSR environment have a greater capacity to enhance filaments’ mechanical properties,
likely due to ECM deposition and stronger cell–matrix interactions. The D filaments also
Figure 6. (a) X-ray diffraction (XRD) of PLA/CPO before and after extrusion, (b) Stress strain diagram
of filaments containing no cells (gray), with KOSR (blue) and with vitamin D, (c) SEM image shows
white areas corresponding to CPO particles and grey areas representing the PLA matrix.
The stress–strain diagram in Figure 6b provides a comparison of the mechanical be-
havior of PLA/CPO composite filaments under different conditions: No Cells (black line),
KOSR (red line), and D (vitamin D-blue line). The figure shows how osteogenic differenti-
ation influences filaments’ mechanical properties. The No Cells control filament, which
was left in solution for 2 weeks without cell integration, exhibits the lowest mechanical
performance, with a tensile strength of approximately 2.8 MPa and a strain of ~0.7%. The
early failure and limited ductility indicate that PLA/CPO degradation weakened the fila-
ment’s structure. On the other hand, PLA/CPO cultured with KOSR and D differentiation
media exhibit a significant increase in both tensile strength and strain, confirming the rein-
forcing effect of cell integration. The KOSR filaments demonstrate the highest mechanical
improvement, with a tensile strength of ~5 MPa and a strain of ~1.6%, outperforming both
the D and the No Cells samples. The almost twofold increase in strength and strain for
both cell-cultured filaments highlights the role of osteogenic differentiation in improving
the samples’ flexibility and structural resilience. These improvements are likely attributed
to early extracellular matrix (ECM) deposition by differentiating cells. As cells adhere to
the scaffold and begin to remodel their microenvironment, secreted collagen and other
matrix proteins form a cohesive interface that enhances the mechanical integrity. Such bio-
logical reinforcement is characteristic of early-stage osteogenesis and reflects the scaffold’s
potential to support functional tissue development.
A key observation is that the KOSR filaments maintain strength over a larger strain
range before ultimate failure. This behavior indicates that osteogenic cells within the KOSR
environment have a greater capacity to enhance filaments’ mechanical properties, likely
due to ECM deposition and stronger cell–matrix interactions. The D filaments also show
J. Manuf. Mater. Process. 2025,9, 149 12 of 14
improved mechanical performance compared to the No Cells control, but its strength and
strain values remain lower than KOSR. Another important distinction is the mode of failure
across the filaments. The No Cells filaments exhibit brittle failure, failing at lower stress and
strain levels, whereas both KOSR and D filaments show improved toughness and resistance
to early failure. This suggests that cellular remodeling not only enhances the load-bearing
capacity but also improves ductility and energy dissipation prior to fracture. This behavior
confirms that biological integration reinforces the filament’s structure, delaying the onset
of failure and allowing greater energy absorption before breaking. These results indicate
that PLA/CPO decomposition over time contributes to mechanical degradation. However,
cell integration significantly strengthens the material, with KOSR treatment yielding the
most pronounced mechanical enhancement.
4. Conclusions
This study enhances bone tissue engineering by introducing a PLA/CPO composite
scaffold fabricated via FDM. The oxygen release of the filaments exhibited a gradual
increase over 3 days, confirming that CPO hydrolysis occurs as the material absorbs fluid.
This early oxygenation may support osteogenic activity and cell viability
in vitro
, which are
essential for bone tissue formation. The research highlights that optimal scaffold quality and
accuracy are achieved at a printing speed of 75 mm/s. The PLA/CPO scaffold, featuring
25% porosity and 0.60 mm pore size, showed enhanced expression of osteogenic markers,
especially when supplemented with osteogenic media, leading to marked upregulation of
the RUNX2 gene which induces the upregulation of SPP1 and SP7, leading to increased
proliferation and differentiation towards osteobalst –lineage. Due to upregulation of
RUNX2, SPP1 and SP7 and relative downregulation of COL1a1 would be indicative of early
Osteoblast-lineage differentiation. Mechanical testing revealed improved properties after
cell culturing, with notable increases in tensile strength (approximately 2 MPa) and strain
(around 0.7%). These results suggest that the PLA/CPO scaffold supports effective bone
regeneration highlighting its potential for further investigation as a scaffold system in bone
tissue engineering.
Author Contributions: Conceptualization, A.M. (Adnan Memi´c), A.T., W.R.W., A.A.M., A.M. (Abdul-
lah Mohammed), M.A., A.E., H.H. and K.E.; methodology, A.M. (Adnan Memi´c), A.T., W.R.W. and
M.A.; software, A.M. (Abdullah Mohammed), A.T., W.R.W. and M.A.; validation, A.M. (Abdullah
Mohammed), A.T., W.R.W., A.M. (Adnan Memi´c), M.A. and H.H.; formal analysis, A.M. (Adnan
Memi´c), A.T., W.R.W., A.A.M., A.M. (Abdullah Mohammed), M.A., A.E., H.H. and K.E; investiga-
tion, A.M. (Abdullah Mohammed), A.T., W.R.W. and M.A.; resources, A.T., W.R.W., A.A.M., A.M.
(Adnan Memi´c), A.E., H.H. and K.E.; data curation, A.M. (Abdullah Mohammed), A.T., W.R.W. and
M.A.; writing—original draft preparation, A.M. (Abdullah Mohammed), A.T., W.R.W., M.A. and
H.H.; writing—review and editing, A.M. (Adnan Memi´c), A.T., W.R.W., A.A.M., A.M. (Abdullah
Mohammed), M.A., A.E., H.H. and K.E.; visualization, A.M. (Abdullah Mohammed), A.T., W.R.W.,
M.A. and H.H.; supervision, W.R.W., A.A.M., A.M. (Adnan Memi´c), A.E., H.H. and K.E.; project
administration, W.R.W., A.A.M., A.M. (Adnan Memi´c), A.E., H.H. and K.E.; funding acquisition,
W.R.W., A.A.M., A.M. (Adnan Memi´c), A.E., H.H. and K.E. All authors have read and agreed to the
published version of the manuscript.
Funding: Deanship of Scientific Research (DSR) at King Abdulaziz University, Jeddah: GPIP:1846-
903-2024.
Data Availability Statement: The data supporting the conclusions of this article will be made
available by the authors on request.
J. Manuf. Mater. Process. 2025,9, 149 13 of 14
Acknowledgments: This project was funded by the Deanship of Scientific Research (DSR) at King
Abdulaziz University, Jeddah, under grant no. (GPIP:1846-903-2024). The authors, therefore, ac-
knowledge with thanks DSR for technical and financial support.
Conflicts of Interest: The authors declare no conflict of interest.
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