Recent Advances in Bone Tissue Engineering: Enhancing the Potential of Mesenchymal Stem Cells for Regenerative Therapies

Abstract

Bone tissue engineering (BTE) has emerged as a promising strategy for addressing bone defects and disorders that cannot be repaired through traditional methods. This field leverages the potential of various biomaterials, cells, and bioactive factors to promote bone regeneration. Mesenchymal stem cells (MSCs) have gained significant attention due to their osteogenic potential, which can be enhanced through osteoinductive factors. Osteoinductive factors, including growth factors like BMPs, TGF-β, VEGF, and IGF, play a crucial role in stimulating the osteodifferentiation process, thereby promoting bone regeneration. Furthermore, bioprinting technologies have opened new avenues for precisely designing scaffolds that can mimic the native bone architecture and provide a conducive environment for MSC differentiation. The integration of bioprinting with mesenchymal stem cells and osteoinductive factors has the potential to revolutionize regenerative therapies by allowing for the creation of patient-specific bone grafts. This review highlights the latest developments in MSC-based therapies, the role of osteoinductive factors, and the impact of bioprinting in advancing BTE. It also discusses future directions for improving the efficacy and clinical translation of these technologies.

Full text

Academic Editor: Madhav Bhatia
Received: 18 March 2025
Revised: 9 April 2025
Accepted: 14 April 2025
Published: 17 April 2025
Citation: Kostadinova, M.;
Raykovska, M.; Simeonov, R.; Lolov, S.;
Mourdjeva, M. Recent Advances in
Bone Tissue Engineering: Enhancing
the Potential of Mesenchymal Stem
Cells for Regenerative Therapies. Curr.
Issues Mol. Biol. 2025,47, 287. https://
doi.org/10.3390/cimb47040287
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/).
Review
Recent Advances in Bone Tissue Engineering: Enhancing the
Potential of Mesenchymal Stem Cells for Regenerative Therapies
Milena Kostadinova 1, * , Miryana Raykovska 2, Radoil Simeonov 3, Stephan Lolov 1
and Milena Mourdjeva 1
1Institute of Biology and Immunology of Reproduction, Department of Molecular Immunology, Bulgarian
Academy of Sciences, 1113 Sofia, Bulgaria; dr_lolov@yahoo.com (S.L.); milena_mourdjeva@abv.bg (M.M.)
2Institute of Information and Communication Technologies, Bulgarian Academy of Sciences,
1113 Sofia, Bulgaria; mirianaraykovska@gmail.com
3Department of Orthopedics and Traumatology, University Hospital Queen Giovanna-ISUL,
1527 Sofia, Bulgaria; radoil_shupi@abv.bg
*Correspondence: mkostadinova@ibir.bas.bg
Abstract: Bone tissue engineering (BTE) has emerged as a promising strategy for addressing
bone defects and disorders that cannot be repaired through traditional methods. This field
leverages the potential of various biomaterials, cells, and bioactive factors to promote bone
regeneration. Mesenchymal stem cells (MSCs) have gained significant attention due to
their osteogenic potential, which can be enhanced through osteoinductive factors. Osteoin-
ductive factors, including growth factors like BMPs, TGF-
β
, VEGF, and IGF, play a crucial
role in stimulating the osteodifferentiation process, thereby promoting bone regeneration.
Furthermore, bioprinting technologies have opened new avenues for precisely designing
scaffolds that can mimic the native bone architecture and provide a conducive environ-
ment for MSC differentiation. The integration of bioprinting with mesenchymal stem cells
and osteoinductive factors has the potential to revolutionize regenerative therapies by
allowing for the creation of patient-specific bone grafts. This review highlights the latest
developments in MSC-based therapies, the role of osteoinductive factors, and the impact of
bioprinting in advancing BTE. It also discusses future directions for improving the efficacy
and clinical translation of these technologies.
Keywords: mesenchymal stem cells; critical-size bone defects; bone tissue engineering;
bioprinting; osteogenic differentiation; osteoinductive factors
1. Introduction
Bone is the second most transplanted tissue after blood; over two million bone grafting
procedures are performed worldwide annually [1,2].
Bones are important organs that perform a wide range of functions, including struc-
tural support for the rest of the body, protection of vital internal organs and structures,
maintenance of mineral homeostasis and acid–base balance, serving as a depot for growth
factors and cytokines, and providing an environment for hematopoiesis in the bone mar-
row [
3
]. They mainly comprise osteocytes, osteoblasts, osteoclasts, and the extracellular
matrix (ECM), which maintains a dynamic balance between bone resorption and bone
formation [
4
,
5
]. Furthermore, bone is a vascularized organ with a unique capacity for
self-repair after damage, although this ability is limited to small fractures with a size of a
few millimeters [
2
]. Large bone defects in adults and children over two years of age do not
reossify successfully, representing a significant biomedical problem [6].
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Curr. Issues Mol. Biol. 2025,47, 287 2 of 28
In the case of large bone defects, treatment with various devices and implants is
necessary to restore the structural functionality of the tissue. In order to support complete
bone regeneration, diverse autografts, allografts, and xenografts are used [1]. Autologous
bone grafts are the gold standard for successfully treating bone defects of critical dimen-
sions. However, damage to the donor site, the limited amount of autologous grafts, and
subsequent complications limit their utility. Allografts are an alternative but also present
problems such as immunological rejection and the possibility of pathogen transmission [
2
].
Another common approach is placing non-biodegradable but mechanically stable implants,
usually based on titanium, stainless steel, or ceramic [
7
]. Nonetheless, this approach al-
most always introduces the need for a secondary operative intervention to remove the
implant from the body due to the occurrence of different post-surgery complications [
7
].
The limitations of traditional bone grafting methods were discussed recently by Hoveidaei
et al. [
8
], all leading to the requirement of developing new approaches, including BTE
and, in particular, bioprinting. At the moment, there is no unified solution, nor is there a
generally accepted experimental model of large bone defects.
Several surgical specialties are related to bone defects. When discussing the means to
overcome critical-size bone defects, all measures are related to three main aspects of the task,
but each specialty emphasizes a different aspect. The main parameters of the “solution”
are (1) appropriate shape/dimensions and mechanical strength; (2) biological tolerance
through the permanent incorporation of or gradual replacement with the patient’s own
tissue; (3) relation to existing or emerging infectious processes. For example, in orthopedics,
the mechanical strength of the implant is of paramount importance. When the greatest
problem is the presence of a cavity (like after a mastoidectomy), efforts are directed towards
achieving volume by transplanting muscle and adipose tissue in addition to bone. In this
regard, initial studies were aimed at replacing only the “intercellular substance”. Especially
in traumatology and otosurgery, the risk of existing or emerging infection is significant.
Therefore, the introduction of effective antimicrobial agents into the structure of the scaffold
or even into its material composition is another important direction of scientific research.
Reconstruction in the dental and maxillofacial region presents a significant challenge
because of the complex nature of the tissues and the functional demands placed upon
them [9,10].
Still, orthopedics is the leading clinical discipline in terms of implant needs. The
most common cause of bone damage is mechanical trauma, but other conditions are also
possible, like infections, tumors, and loosening of endoprostheses of large joints. In multi-
fragment fractures with avital fragments, or missing ones, bone loss is present even before
the regeneration process. Bone and joint infections, whether primary or iatrogenic, are
another common cause of bone loss. As a result of bacterial metabolism, which releases
many exo- and endotoxins, many of which are lytic enzymes, bone resorption is observed.
Tumors are undoubtedly one of the most common and severe causes of bone loss. In the
radical removal of bone neoplasias, a wide resection of the affected bone is required, the
replacement of which is seriously challenging. All these reasons lead to serious disability
in patients and to long periods of treatment, often unsuccessful.
In addition to the problems described above, an increasing number of people suffer
from osteoporosis due to the aging of the population. This reduces the quality of bone
tissue and adversely affects the treatment of emerging bone damage [
11
], widening the
group of people inquiring about bone substitution.
2. Bone Tissue Engineering
Personalized medicine in bone regeneration is based on an approach tailored to the
specifics of the individual patient so that the grafts applied perfectly match the shape,

Curr. Issues Mol. Biol. 2025,47, 287 3 of 28
structure, and dimensions of the defect site. In addition, cells isolated from the patient can
be used and further integrated into this personalized implant, representing an autologous
strategy that reduces the risk of immune rejection and inflammation [12,13].
In BTE, three key factors are required for successful bone regeneration: (1) osteopro-
genitor cells, such as embryonic stem cells (ESCs), mesenchymal stem cells (MSCs), or
induced pluripotent stem cells (iPSCs), which are capable of differentiating into functional
bone cells; (2) specific growth factors that stimulate cell migration, proliferation, differentia-
tion, and vascularization; and (3) biomaterials that offer a three-dimensional matrix for cell
adhesion and growth.
Several methods of scaffold creation are widely used nowadays. Additive manufac-
turing (AM) techniques, such as fused filament fabrication (FFF) and 3D printing, have
been used to generate BTE scaffolds, offering advantages in controlling scaffold structural
properties such as pore size, degree of porosity, and mechanical strength [
13
,
14
]. FFF is
the basic additive manufacturing process, in which the source material (a thermoformable
polymer) is heated to a viscous state and extruded through a nozzle that moves along a set
two-dimensional (2D) tool path, layer by layer [
15
]. AM techniques can be successfully im-
plemented in customized BTE by generating a 3D computer-assisted design (CAD) model
of the individual patient’s anatomical structure [
14
,
16
]. Various synthetic polymers, such as
polycaprolactone (PCL), poly(lactic-co-glycolic) acid, and polyethylene glycol [
17
], natural
polymers such as chitosan and cellulose, and also natural bone substitutes, such as calcium
sulfate [
2
], tricalcium phosphate (TCP) [
18
,
19
], ceramics [
20
], bioactive glass [
21
], and
calcium phosphate cement [
22
], are used in BTE and are necessary for bone regeneration to
provide osteoconductive support for space filling and mechanical stability. However, their
lack of/light osteoinductivity remains a challenge in their use as an alternative to natural
bone [2].
Bone substitutes based on natural polymers are considered next-generation grafts.
The use of
β
-tricalcium phosphate (
β
-TCP) and hydroxyapatite (HA) with biopolymers
such as collagen and alginate has been studied extensively [2,23–26].
The three most commonly used mineral supplement forms of calcium phosphate [
27
]
are TCP [
28
], HA crystals [
29
], and decellularized bone matrix (DCB) [
30
]. They all differ in
form and function:
TCP contains readily available calcium and phosphates for bone production and is
degraded into ions via hydrolysis and via osteoclast resorption [31,32].
HA is the naturally occurring crystalline form of bone mineral and is similar to TCP,
except for its denser crystal structure and increased mechanical properties [27,33].
DCB is derived from natural (xenogeneic and allogeneic) bone sources and may
include an organic protein phase such as collagen. However, a clinically available form
of DCB, Bio-Oss, provides millimeter-size granules of bovine trabecular bone, with the
organic phase largely removed [27,33].
An essential aspect of BTE is that the engineered and fabricated bone support is
gradually degraded and replaced by new bone tissue without adverse health effects. It is
important that the scaffold contributes to bone regeneration by providing a suitable envi-
ronment for stem cells with osteogenic potential to differentiate and regenerate new bone.
Another essential requirement for bone regeneration scaffolds is optimal vascularization.
In the process of defect regeneration, blood vessels development precedes bone formation
and only after functional circulation provides nutrients can new bone formation occur [
34
].
As mentioned above, the inflammation environment often accompanies critical-size bone
defects and should be considered when designing the scaffold properties, so the inclusion
of antimicrobial/antiinflammation material additives is of critical importance.

Curr. Issues Mol. Biol. 2025,47, 287 4 of 28
Scaffolds used for tissue engineering must degrade at an appropriate rate for tissue
regeneration while mechanically supporting the defect area. It is known that the regenera-
tion of bone defects of critical dimensions requires time, and therefore, the scaffolds must
have an adequate degradation rate, providing the necessary structural support to the tissue
during the repair process [
1
,
35
]. On the other hand, the cells in the scaffold play an equally
important role in tissue remodeling and must be able to interact with the surrounding
tissue easily. Figure 1summarizes the requirements that scaffolds designed and created for
bone regeneration should meet.
Curr. Issues Mol. Biol. 2025, 47, 287 4 of 29
bone defects and should be considered when designing the scaffold properties, so the in-
clusion of antimicrobial/antiinflammation material additives is of critical importance.
Scaffolds used for tissue engineering must degrade at an appropriate rate for tissue
regeneration while mechanically supporting the defect area. It is known that the regener-
ation of bone defects of critical dimensions requires time, and therefore, the scaffolds must
have an adequate degradation rate, providing the necessary structural support to the tis-
sue during the repair process [1,35]. On the other hand, the cells in the scaffold play an
equally important role in tissue remodeling and must be able to interact with the sur-
rounding tissue easily. Figure 1 summarizes the requirements that scaffolds designed and
created for bone regeneration should meet.
Figure 1. Scaffolds designed for bone regeneration must meet various requirements in terms of
structural and biological properties.
As a mandatory step in the creation of a personalized bone implant, the construction
of a matrix from a suitable material is relatively well studied. It should be noted that some
materials have already found real clinical applications; these include polymethylmethac-
rylate for calvaria defects [36], PCL and β-TCP for the facial skull [37], and titanium for
the femur [38]. Additional sources of information are clinical studies of application in non-
osseous structures, as well as numerous animal models [39–41].
Bone cells are extremely sensitive to the chemical and physical properties of the scaf-
folds on which they are cultured. Their material composition, roughness, and topography
contribute to the osteogenic process, determining cell growth and differentiation. Cell ad-
hesion to the scaffold material represents the initial phase of cell–scaffold communication,
triggering multiple cellular responses, including proliferation and differentiation, as well
as vascularization [33,42].
Figure 1. Scaffolds designed for bone regeneration must meet various requirements in terms of
structural and biological properties.
As a mandatory step in the creation of a personalized bone implant, the construction
of a matrix from a suitable material is relatively well studied. It should be noted that
some materials have already found real clinical applications; these include polymethyl-
methacrylate for calvaria defects [
36
], PCL and
β
-TCP for the facial skull [
37
], and titanium
for the femur [
38
]. Additional sources of information are clinical studies of application in
non-osseous structures, as well as numerous animal models [39–41].
Bone cells are extremely sensitive to the chemical and physical properties of the scaf-
folds on which they are cultured. Their material composition, roughness, and topography
contribute to the osteogenic process, determining cell growth and differentiation. Cell
adhesion to the scaffold material represents the initial phase of cell–scaffold communication,
triggering multiple cellular responses, including proliferation and differentiation, as well
as vascularization [33,42].
3. Bioprinting
Bioprinting is a rapidly developing multidisciplinary technology, the goal of which
is to produce vital three-dimensional constructs by depositing a biocompatible material
(bioink) containing living cells by printing layer by layer on patient-specific digital mod-

Curr. Issues Mol. Biol. 2025,47, 287 5 of 28
els [
8
,
43
,
44
]. Bioprinting has already enabled the production of small units of tissues [
45
,
46
]
and organoids [
47
] that share some of the functions of their native counterparts [
48
] and
are useful as
in vitro
models for research and as test platforms for screening or developing
drugs and personalized therapies [
47
]. Furthermore, 3D bioprinting has been used to pro-
duce various human tissues such as skin [
49
], heart [
50
], bone [
51
], liver [
52
], and nerve [
53
]
tissues, as well as the extracellular matrix [
54
,
55
]. The advancement of bioprinting will
allow patients to obtain access to 3D-printed tissues and organs that can replace lost or
damaged ones [56].
Today, the most reported bioprinting technologies are extrusion, injected/droplet
bioprinting, laser-assisted bioprinting (LAB), and stereolithography [43,57].
Extrusion-based bioprinting is the simplest and most widely used form, in which the
bioink is ejected/expelled by a pneumatic, piston, microfluidic, or screw-based filament
deposition mechanism [
58
]. Extrusion technology has a wide range of biocompatible
materials that can be printed (e.g., cell-loaded hydrogels), but its accuracy is usually
limited [
59
]. The viscosity and density of the bioink, the liquid phase of the bioink, the
extrusion speed, and other material-specific properties, such as the ability to crosslink
between the printed layers, are some of the main factors that need to be considered to
achieve quality. The main disadvantages of extrusion bioprinting technology include low
printable resolution, combined with possible nozzle clogging and shear stress-induced cell
damage during extrusion [60].
Injected bioprinting, also referred to as droplet bioprinting, deposits droplets of
bioink under the control of a piezoelectric or thermal system [
61
]. The process is performed
continuously to form CAD patterns layer by layer. Droplet bioprinting offers the advantages
of relatively low cost, high accuracy and speed, compatibility with various biological
materials, and also the ability to deposit multiple cell types simultaneously [
60
]. However,
it is not possible to print high cell concentrations, and low-viscosity materials reduce
structural strength [
62
]. Another disadvantage is the need for support material, as well as
possible nozzle clogging [60].
In LAB, cells are printed onto a receiving substrate by a pulsed laser beam at controlled
speeds [
63
]. The receiving substrate has a ribbon structure that consists of an energy-
absorbing layer on top (e.g., titanium or gold) and a bioink layer (e.g., cells and hydrogel)
at the bottom. During printing, focused pulses from the laser source stimulate a region of
the energy-absorbing layer. The energy vaporizes part of the donor layer to create a high-
pressure bubble that propels the bioink onto the receiving substrate in the form of droplets.
So, LAB is non-contact printing and avoids the problem of nozzle clogging in extrusion
or droplet bioprinting, allowing for printing with higher cell densities. Furthermore,
LAB does not create mechanical stress on the cells during printing, resulting in high
cell viability [
64
]. Recent studies have demonstrated the benefits of prevascularization
using LAB to promote vascularization and bone regeneration. In the work of Kérourédan
et al. [
65
], the preorganization of human umbilical vein endothelial cells onto a newly
developed biopaper using LAB led to the formation of microvascular networks. However,
the printing cost is higher, and the throughput is lower [
44
,
62
,
66
]; there are high crosslinking
requirements, and the effect of the laser on cells is still unclear [64].
Stereolithography is a nozzleless system that uses light to polymerize photosensitive
inks in a layer-by-layer deposition process. Stereolithography is a relatively low-cost,
short-time printing method and results in high resolution. Inks with high viscosity can be
printed, and great cell viability can be achieved despite the possible damaging effect of UV
light in photocuring. However, the reduced selection of photosensitive biomaterials limits
its use [
57
]. Other disadvantages are the lack of multi-cell-type printability, the need for
support materials, and expensive equipment [60].

Curr. Issues Mol. Biol. 2025,47, 287 6 of 28
Table 1lists the most often reported bioprinting techniques in recent years, organized
by major characteristics based on studies by Mobaraki et al. [44] and Sousa et al. [60].
Table 1. Comparison of bioprinting methods in terms of their major characteristics. The “+” symbol
indicates the relative performance level for each characteristic, where “++++” means excellent, and
“+” means low level.
Resolution Speed Mechanical Properties Porosity Cost
Extrusion
bioprinting + ++ ++ + ++
Injected bioprinting ++ ++++ + ++ +
Stereolithography +++ + ++++ +++ +++
LAB ++++ +++ +++ ++++ ++++
Special attention should be paid to the macro- and microstructure of the print. The
macrostructure is of particular importance for the accurate filling of the bone defect. Creat-
ing an implant with smaller dimensions implies a slightly longer healing process but easier
intraoperative positioning. Generating an implant that is larger than the defect requires
intraoperative fitting, probably with a better early mechanical result.
The microstructure of the bioprint is critical when it comes to parenchymal organs.
Both the size and arrangement of different cell types must be taken into account. Factors
such as resolution, structural fidelity, and hierarchical architectures have been extensively
discussed [
67
,
68
]. Specifically for supporting tissues, and bones in particular, several
factors should be considered: the correspondence between the thickness and orientation
of the newly constructed lamellae (in view of the targeted mechanical resistance), the
distance between them (for easier cell population and metabolism), and possibly their
spatial orientation (as a prerequisite for neovascularization). The mechanisms of bone
remodeling under compressive and tensile forces are relied upon to optimize construction
at a later stage, while cellular self-organization during co-cultivation is utilized for the
proper formation of new vessels [69].
Regardless of the method, the functioning of any of these bioprinting platforms re-
quires bioinks with a specific set of properties, such as chemical composition, viscosity, and
material rheology. In fact, the bioink is a key factor in the bioprinting strategy and is the
defining element that distinguishes “bioprinting” from “3D printing”. The main advantage
of bioprinting over 3D-printed cell-free scaffolds that are loaded with cells in post-printing
modification procedures is that bioprints incorporate osteoinductive stem cells and os-
teoinductive factors, creating a microenvironment that enhances the regeneration of the
bone [69].
4. Bioink
The success of generating functional tissues largely depends on the qualities of the
bioink [
43
]. It must exhibit appropriate mechanical properties, stability, and biological
activity to achieve and maintain the designed structures while providing a favorable envi-
ronment for the cell population. The most common materials used are natural polymers [
44
],
but there are also bioinks based on synthetic polymers [
64
] and hybrid/composite bioinks.
Polymeric bioinks are preferred for their low cost, biocompatibility, biodegradability,
and safe processing. The main natural polymers used in bioprinting [
44
] are collagen [
47
,
70
],
fibrin [
71
], silk [
72
], chitosan [
73
], alginate [
74
], hyaluronic acid [
75
],
ε
-polylysine [
76
], and
gelatine methacrylate [
76
,
77
]. Water-soluble polymers, known as hydrogels, are the most
used bioinks due to their chemical configuration and suitable conditions for cell growth in

Curr. Issues Mol. Biol. 2025,47, 287 7 of 28
3D [
76
,
78
]. Hydrogels are three-dimensional polymer networks composed primarily of pro-
teins, peptides, and polysaccharides; they are biocompatible materials that can retain a large
amount of water. As a result, they provide a favorable environment for cells [
79
], as they
are characterized by excellent permeability for nutrients, oxygen, and other water-soluble
compounds [
80
], have a high capacity for drug delivery, biocompatibility, and sensitivity
to environmental stimuli such as temperature, pH, and solvent type [
76
]. Polysaccharide-
based hydrogels comprise chitosan, alginates, agarose, hyaluronic acid, glucans, cellulose,
guar gum, and cyclodextrin polymers. Natural polymer hydrogels offer numerous advan-
tages, including abundant sources, porous structures, multifunctional groups, favorable
swelling properties, biodegradability, and low immunogenicity [76].
In contrast, synthetic materials have various advantages over natural materials, such
as the capacity to be tailored with specific physical properties and increased uniformity. The
following are some examples of synthetic materials used in this field: poly (
ϵ
-caprolactone),
poly(lactic-co-glycolic) acid, poly(l-lactic) acid, polyglycolic acid, polyurethane, polyethy-
lene glycol, polyether ether ketone, polyvinylpyrrolidone, and Pluronic. Nevertheless,
synthetic materials for 3D bioprinting have some disadvantages, such as poor biocompat-
ibility, the potential release of toxic degradation products, and the absence of bioactive
ligands [60].
Composite bioinks, also known as hybrid bioinks, are blends or multi-phase materials
composed of two or more polymers with integrated bioactive inorganic additives and
encapsulated cells [
81
]. Although polymeric materials for 3D-bioprinted structures have
many advantages (low weight, low melting point, lower cost, and processing flexibility), the
main problem is their low mechanical strength and functionality. To overcome this obstacle,
various additives are added to the polymer matrix. One of the most used ceramic materials
in tissue engineering applications is HA. Its powder is widely used in 3D printing [
82
].
Other used nanocomponents are
β
-TCP [
83
] and clay [
84
], metallic nanoparticles [
85
], and
nano-carbon in the form of graphene [86] and carbon nanotubes [44,87].
Bioinks require crosslinking (gelling or hardening), and the shear forces resulting from
this process can cause damage to the cells.
The quality of the final print product and its cell content depends on the rheological
properties and crosslinking mechanisms of the bioink. During the printing process and
depending on the printing technology used, several biomechanical parameters that affect
the bioink must be considered for achieving the generation of an appropriate tissue ar-
chitecture with good form integrity. These include flow pattern, viscosity, viscoelasticity,
surface tension, flow rate, and mechanical forces such as hydrostatic pressure, shear stress,
and extensional stress, which will define the printability of a biomaterial [44,88,89].
Shear stress is a mechanical force that occurs when fluids (e.g., bioinks) move over
solid surfaces or through narrow channels [
90
]. In extrusion and injected bioprinting
techniques, bioinks flow through the printer nozzle, where the associated shear forces can
induce cellular damage [
57
,
91
]. For MSCs, excessive shear stress can lead to increased het-
erogeneity of cell response, induce inelastic ultra-structural distortion of the cell membrane
and chromatin, and increase necrotic subpopulations post-printing [
92
]. The results of a
study by Lemarié et al. showed that the higher the shear stress values, the lower the viable
cell populations. Regarding the quantification of the fibroblasts damaged (lysis, necrosis,
and apoptosis) during microextrusion, lysis was shown to be the prevailing cell death
pathway compared to necrosis and apoptosis [
90
]. However, there is evidence that shear
stress is related to stem cells’ fate and can positively influence the osteogenic differentiation
of MSCs [93].
Several approaches can be used to reduce the negative impacts of shear stress on
cells during bioprinting. First of all, applying low-viscosity bioinks reduces the shear

Curr. Issues Mol. Biol. 2025,47, 287 8 of 28
stress experienced by cells during extrusion [
94
]. Another option is the development of
shear-thinning hydrogels [
57
]. Shear-thinning materials are bioinks that exhibit a decrease
in viscosity under shear forces and recover their original viscosity when shear stress is
removed [
95
]. The most logical approach, but one that does not always work, is flow rate
control; the careful modulation of the print head’s flow rate allows for a reduction in shear
forces while maintaining adequate extrusion pressure.
Extrusion pressure is another key factor in the bioprinting process that influences
MSC viability and differentiation. High extrusion pressure can cause several harmful
effects, including the mechanical deformation of MSCs, leading to disrupting cell–cell and
cell–matrix interactions, which may negatively affect their viability [96].
To minimize the negative effects of extrusion pressure on cell content, several strategies
can be employed. By controlling the rate of extrusion pressure, it is possible to minimize the
forces that affect cell survival and differentiation while maintaining the structural integrity
of the printed construct [
97
,
98
]. Another commonly used technique is printing temperature
control: lowering the printing temperature can help to reduce the viscosity of bioinks,
allowing for printing at a lower extrusion pressure [
98
]. A strategy that is not so easy to
implement is print head design optimization: printheads with larger nozzle diameters [
99
]
or multi-channel nozzles [
100
] can help distribute the pressure more evenly, reducing the
mechanical stress on the cells during extrusion.
Crosslinking is a process by which bioinks undergo a chemical or physical transfor-
mation to form a stable network, allowing the printed structure to hold its shape. Common
crosslinking methods include chemical, physical, and thermal crosslinking [
101
]. Chemical
crosslinking involves the use of crosslinking agents to chemically bond the polymer chains
in the bioink. While chemical crosslinking stabilizes the scaffold, it can also have cytotoxic
effects on cells if it is not carefully controlled [
102
]. Physical crosslinking is a method
based on UV light exposure (photoinitiated) or ionic crosslinking (using calcium salts, for
example). These methods are generally less cytotoxic but may still affect MSC viability
depending on the duration and intensity of exposure [
103
]. Thermal crosslinking is a
heat-induced gelation process, but it can cause thermal damage to sensitive cells like MSCs
if temperatures exceed safe thresholds [
104
]. Combining different crosslinking methods,
so-called dual crosslinking, can help optimize the mechanical properties of the scaffold
while minimizing toxicity to MSCs [103].
The rheology of bioinks should be tailored to the specific printing technique and
scaffold design. A shear-thinning rheology is considered an essential property for bioinks,
since it performs two functions. First, the high viscosity of the material at rest ensures the
mechanical stability of the prints. Second, shear-thinning qualities allow the material to
be printed at a substantially lower pressure while maintaining the same printing speed,
reducing the overall hydrodynamic stresses acting on cells inside the nozzles [
94
]. For
example, the use of shear-thinning bioinks that transition from a gel-like state to a more
fluid state under shear stress helps reduce the impact of extrusion forces while maintaining
structural integrity after printing [57].
Bioinks commonly used in BTE are mostly hybrid since they are composed of natural or
synthetic polymers, or a mix of the two, containing metallic or ceramic components, which
enable the creation of structures with strong mechanical properties and simultaneously
release regenerative signals to osteogenic cells to promote further regeneration and repair.
An essential factor to consider in terms of the cellular component of the bioink is the
initial cell concentration. It is important to retain a sufficient number of live cells after the
printing process to achieve the appropriate cell density. Bioprinting, by its mechanism, is
associated with increased pressure [
91
], and in combination with the different crosslinking
options (chemical, UV rays, etc.), it subjects the cells to shear stress. It is known that

Curr. Issues Mol. Biol. 2025,47, 287 9 of 28
fluid shear stress or extensional flow-mediated cell deformations can damage the plasma
membrane and thus lead to reduced survival [
105
]. This problem is even worse in the case
of using hydrogels with high polymer content as the base for the bioink due to the high
forces and stress on the encapsulated cells. Bioprinting of large-scale tissues or organs is
even more challenging in terms of cell viability, as it needs to find a solution to maintain
sufficient cell viability in the first printed layers of the construct when the printing time
is more than a few hours. Cell density also affects the viscosity of the bioink—the higher
the cell density, the higher the average viscosity, and this can affect the fidelity of the final
construct and the functionality of the printed tissue or organ [64,106].
Despite intensive research in this rapidly developing field, data on actual bioprinting
with MSCs, particularly in the area of BTE, remain very limited. More data are needed to
determine the optimal combination of bioinks, cells, and osteoinductive factors, as well as
the technology and parameters of printing, to achieve a sufficiently stable, durable, and
functional replacement for the missing bone.
5. MSCs as Building Blocks for BTE
The initial selection of cells is critical for the successful fabrication of tissues and organs
through bioprinting. They must have specific biological characteristics and functions so
they can proliferate and differentiate into the required final cell types.
In terms of cell sources, MSCs have been widely used for bioprinting and BTE due
to their unique ability to differentiate into multiple cell types, including osteoblasts, and
to secrete various cytokines to regenerate damaged or injured bone tissues [
107
–
109
]. The
majority of research has focused on bone marrow mesenchymal stem cells (BM-MSCs) due
to their high osteogenic potential [
110
] compared to MSCs derived from other sources like
adipose tissue (AT-MSCs) or perinatal tissues.
MSCs have been extensively studied in recent decades due to their functional prop-
erties. First, these cells can be induced to differentiate into any cell type depending on
culture conditions and environmental cues
in vitro
. However, this is just one of the possible
mechanisms of MSCs’ beneficial actions during repair or injury. MSCs’ biological properties
put them in the focus of interest as building blocks for regenerative medicine because of
their ability to migrate to the site of injury and serve as a conductor of the restoration of
homeostasis, synchronizing the production of cells to be differentiated, secreting growth
and angiogenic factors, and guiding the inflammatory status of the regenerating tissue,
switching between a pro- and anti-inflammatory phenotype. Thus, MSCs loaded into
bioinks for BTE contribute to multiple levels of bone regeneration.
For more than three decades, MSCs have been an important research interest. They
were first isolated from bone marrow, and later, MSCs were shown to reside in a perivascu-
lar niche
in vivo
[
111
]. Their localization near the blood vessels explains why MSCs are
nearly ubiquitous and can be isolated from most tissues—virtually all vascularized organs,
including adipose tissue, articular cartilage, the brain, dental tissues, the endometrium and
menstrual blood, the skin, and perinatal organs and tissues, including amniotic fluid, the
amniotic membrane, the placenta, Wharton’s jelly, umbilical cord tissue, and cord blood. To
identify MSCs, the International Society of Cellular Therapy postulated four minimum cri-
teria, namely, (1) fibroblast-like morphology, (2) plastic adherence, (3) trilineage capacity for
differentiation into osteoblasts, adipocytes, and chondrocytes, and (4) the expression of cell
surface proteins CD73, CD90, and CD105 while lacking the expression of lineage-specific
markers CD45, CD34, CD14, CD19, CD11b, and HLA-DR [
112
]. Although this has been
the main standard for almost three decades, Pittenger et al. recently discussed the need
for genome-wide gene expression studies to provide insights into the biological nature of
MSCs, their expected physiological function, their role in disease pathophysiology, and

Curr. Issues Mol. Biol. 2025,47, 287 10 of 28
their probable therapeutic mode of action [
109
]. These characteristics have not yet been
widely introduced into the laboratory practice of groups working with MSCs, but it is good
to take them into account because such studies can best characterize MSCs’ composition
and function prior to their administration to patients. Promising results from clinical
trials led to the FDA approval of MSC-based therapies, and ongoing investigations are
exploring their potential in cartilage and bone repair, wound healing, and autoimmune
disorders [113,114].
MSCs are considered the most reliable source of osteoprogenitor cells [
109
,
115
] and
play a significant role in the initial formation and maintenance of bone. Endochondral
ossification is a bone-healing mechanism that begins with MSCs that differentiate into
chondrocytes to form cartilage, which is then calcified and subsequently remodeled into
bone [
116
]. Intramembranous ossification is another bone regeneration mechanism in-
volving MSCs and undifferentiated bone progenitors that directly differentiate into os-
teoblasts [
117
]. The plethora of bioactive molecules released by MSCs actively helps create
an optimal regenerative microenvironment [118,119].
The specificities of
in vitro
osteogenic differentiation were recently summarized by
Romano et al. [
120
], and they are as follows: (1) osteogenic induction is usually obtained
in about 21 days by adding dexamethasone,
β
-glycerophosphate, and ascorbic acid to
basic growth media; (2) specific histological staining, which reveals calcium deposits and
the mineralized matrix, verifies the osteogenic phenotype (Von Kossa or Alizarin Red);
(3) an increased expression of bone markers, such as Runt-related transcription factor 2
(Runx2), markers of osteoblast differentiation, such as osteonectin and osteocalcin, as well
as increased synthesis of alkaline phosphatase (ALP) and collagen type I can be detected at
both the mRNA or protein levels.
To elucidate the regulatory mechanism of MSC differentiation in osteoblast cells, Chen
et al. [
121
] used single-cell multiome data to identify the regulatory elements and predict
the regulation relationships between the genes and key transcription factors involved in os-
teogenic differentiation. The results revealed that common regulatory networks among the
four donors’ MSCs were related to the stemness function of the cells, while donor-specific
ones were related to their differentiation into specific cell types. The results further suggest
that regulatory networks in cells with a higher potential for osteogenic differentiation may
be associated with bone density-related diseases and should be considered in choosing
donor MSCs for regenerative therapies. This can also be one of the explanations for the
heterogeneous results reported in clinical trials with MSCs.
Last but not least, MSCs could have a significant contribution to vascularization
during new bone formation. Still, one of the major challenges in BTE is the induction of
angiogenesis and vasculogenesis in the implanted tissues to supply the newly formed
bones with sufficient nutrients and oxygen and to withdraw the waste metabolites, thereby
supporting tissue growth and maturation [
8
,
47
,
122
]. Transplanted MSCs can contribute
to bone regeneration through angiogenesis stimulation by the expression of angiogenic
factors such as VEGF, TGF-β, SDF-1, and stem cell factor (SCF) [8,123,124].
One of the most commonly used methods for angiogenesis induction in BTE is using a
sustained release of angiogenic growth factors [
8
]. Proper dosing and kinetic management
of angiogenic factors and their interaction with osteogenic processes need to be consid-
ered. Another approach is the bioprinting of stimuli-responsive biomaterials (so-called
4D printing) [
125
]; thus, the need to create vascular-like networks in scaffolds can be
by-passed [
126
]. To overcome the need for a sustained production and release of growth
factors, stem cells can be transfected with growth factor genes [
8
,
120
]. In some studies,
the
in vivo
prevascularization of scaffolds by implantation into a highly perfused tissue,
e.g., subcutaneous or muscular pockets, is shown [
127
]; thus, vessels can grow from the

Curr. Issues Mol. Biol. 2025,47, 287 11 of 28
outside into the center of the material until complete vascularization is achieved. Recently,
Kérourédan et al. reported the development of a novel biopaper able to support prevascu-
larization organized by LAB for bone tissue engineering applications. Gelatin-based sheets
incorporating bioactive glasses (BGs) were produced using various freezing methods and
crosslinking parameters [65].
Co-culturing stem cells or osteoblasts with endothelial cells is a common approach to
enhance vascularization [
8
]. Co-culture studies have demonstrated that there is crosstalk
between endothelial cells and MSCs that can lead to synergistic effects on tissue regen-
eration [
128
]. In a rabbit large segmental bone defect model, the co-culture of BM-MSCs
and endothelial progenitor cells (EPCs) on calcium phosphate ceramic scaffolds for bet-
ter vascularization, and thus osteogenesis, was examined. Different ratios of co-cultures
were first studied
in vitro
, and the results indicated that optimal neovascularization and
osteogenesis occurred at a BM-MSC-to-EPC ratio of 1:3. Enhanced osteoid formation and
bone remodeling, sustained by neovascularization, were observed compared to regular
cultures. But still, suboptimal efficacy was noted compared to autologous bone grafts [
129
].
There are several clinical trials at different phases (I, II, or III) for bone fracture repair
using BM-MSCs, AT-MSCs, umbilical cord-derived MSCs (UC-MSCs), and human amniotic
epithelial cells (ClinicalTrials.gov) which were implanted either via direct injection or after
seeding them onto an osteogenic matrix [123].
MSCs have been suggested to contribute to bone healing through three different
mechanisms: differentiation and replacement [
119
], secretion of cytokines and extracel-
lular vesicles [
107
,
130
], and immunomodulatory activity [
131
,
132
]. It is still difficult to
tell which one is the leading mechanism through which MSCs enhance bone regeneration.
Nevertheless, the enhancement of MSCs’ osteogenesis is directly related to improving the
therapeutic effect of MSC-based BTE [
4
]. There is evidence that MSCs tend to differentiate
into pre-osteoblasts at an intermediate stage instead of directly differentiating into osteo-
cytes. Pre-osteoblasts develop into mature osteoblasts, which synthesize bone matrix and
are then incorporated into the matrix as osteocytes [133].
6. Osteoinductive Factors and Their Application in BTE
Bone regeneration is a complex phenomenon that requires active signaling by numer-
ous biomolecules at different stages of bone development [
2
], and signaling pathways, such
as signaling through transforming growth factor
β
(TGF-
β
), Bone Morphogenic Proteins
(BMPs), Wnt, and SHH, regulate the whole process. The targets of these signaling pathways
are Runx2 and osterix (OSX), key transcription factors in the process of MSCs’ osteogenic
differentiation [
4
]. Upon specific signals, the MSCs that reside in the bone marrow and
periosteum differentiate into osteoblasts [
134
–
136
]. MSCs’ commitment to the osteoblastic
phenotype is driven by a series of transcription factors in which Runx2/core-binding factor
subunit alpha-1 (Runx2/Cbfa1) and the OSX are crucial for osteoblast differentiation, and
the depletion of either of these two factors provokes complete damage to the mineralized
skeleton [
137
]. When MSCs enter the pre-osteoblastic commitment pathway, a proliferative
phase is activated, together with a high expression of ALP, an early osteogenic marker
and critical for subsequent bone mineralization [
138
]. The transition to mature osteoblasts
is marked by the expression of osteocalcin (OCN), bone sialoprotein (BSP), osteopontin
(OPN), and collagen-I, which further contribute to the formation of the osteoid structure
and mineralization [136,139].
Because of the essential role of growth factors in controlling cellular processes [
140
]
and their ability to directly promote tissue regeneration, a wide range has been investigated
and tested for therapeutic applications [
141
], including bone regeneration [
142
,
143
]. Factors
such as fibroblast growth factor (FGF), vascular endothelial growth factor (VEGF), insulin-

Curr. Issues Mol. Biol. 2025,47, 287 12 of 28
like growth factor (IGF), TGF-
β
, platelet growth factor (PDGF), and BMPs are the main
participants in the process of bone regeneration [144].
A complex spatiotemporal cascade of cytokines orchestrates healing after bone frac-
ture [
145
,
146
] (Figure 2). Inflammatory cytokines such as IL-1
β
, IL-6, IL-17, and tumor
necrosis factor
α
(TNF-
α
) induce invasion by lymphocytes, plasma cells, macrophages,
and osteoclasts [
147
]. Invading macrophages clear the necrotic centers and release TNF
α
,
stimulating increased osteoclast activity. Osteoclasts resorb the fragmented bone mass,
releasing incorporated IGF and BMPs, which induce the osteoblastic differentiation of
osteoprogenitor cells or MSCs [
146
,
148
,
149
]. Neovascularization at the fracture site occurs
early in this process as endothelial cells induce angiogenesis in response to VEGF and
low oxygen concentration [
146
,
150
]. Endothelial cells are the primary source of BMPs at
the fracture site, leading to the osteogenesis of osteoblasts. Osteoid production by these
osteoblasts begins outside the fracture, creating a callus and mechanically integrating the
bone [
146
]. PDGF, TGF-
β
and FGF, released by plasma cells, macrophages, and osteoblasts,
induce and maintain cell proliferation and differentiation [149,151,152].
Curr. Issues Mol. Biol. 2025, 47, 287 12 of 29
A complex spatiotemporal cascade of cytokines orchestrates healing after bone
fracture [145,146] (Figure 2). Inflammatory cytokines such as IL-1β, IL-6, IL-17, and tumor
necrosis factor α (TNF-α) induce invasion by lymphocytes, plasma cells, macrophages,
and osteoclasts [147]. Invading macrophages clear the necrotic centers and release TNFα,
stimulating increased osteoclast activity. Osteoclasts resorb the fragmented bone mass,
releasing incorporated IGF and BMPs, which induce the osteoblastic differentiation of
osteoprogenitor cells or MSCs [146,148,149]. Neovascularization at the fracture site occurs
early in this process as endothelial cells induce angiogenesis in response to VEGF and low
oxygen concentration [146,150]. Endothelial cells are the primary source of BMPs at the
fracture site, leading to the osteogenesis of osteoblasts. Osteoid production by these
osteoblasts begins outside the fracture, creating a callus and mechanically integrating the
bone [146]. PDGF, TGF-β and FGF, released by plasma cells, macrophages, and
osteoblasts, induce and maintain cell proliferation and differentiation [149,151,152].
Figure 2. Osteoinductive factors and cytokines involved in bone tissue regeneration of a critical-
sized bone defect after injury/trauma. Stages, basic processes, and cells responsible for the
regeneration process.
The inclusion of factors in bioinks is an approach to induce proliferation and
differentiation in a time-dependent manner [8]. Additionally, scaffolds containing growth
factors provide a beer option for personalized treatment, bone defect repair, and bone
regeneration in orthopedics [34]. In BTE, various approaches, such as physical
entrapment, chemical binding, surface modifications, biomineralization, micro- and
nanoparticle encapsulation, and genetically modified cells, are applied to provide
spatiotemporal control of the delivery of a given growth factor from the scaffold [146]. Some
key growth factors described above as critical during normal bone healing have been used in
clinical approaches to treat bone nonunions. The timing of therapeutic growth factor delivery
is crucial to optimize tissue induction while minimizing unwanted or inhibitory effects. A
disadvantage of this strategy is the short half-life and high clearance rate of growth factors in
vivo, especially when administered systemically [142,153]. The incorporation of growth
factors into scaffolds for BTE has significant potential to improve therapeutic outcomes [2,146].
Figure 2. Osteoinductive factors and cytokines involved in bone tissue regeneration of a critical-sized
bone defect after injury/trauma. Stages, basic processes, and cells responsible for the regeneration
process.
The inclusion of factors in bioinks is an approach to induce proliferation and differen-
tiation in a time-dependent manner [
8
]. Additionally, scaffolds containing growth factors
provide a better option for personalized treatment, bone defect repair, and bone regenera-
tion in orthopedics [
34
]. In BTE, various approaches, such as physical entrapment, chemical
binding, surface modifications, biomineralization, micro- and nanoparticle encapsulation,
and genetically modified cells, are applied to provide spatiotemporal control of the delivery
of a given growth factor from the scaffold [
146
]. Some key growth factors described above
as critical during normal bone healing have been used in clinical approaches to treat bone
nonunions. The timing of therapeutic growth factor delivery is crucial to optimize tissue
induction while minimizing unwanted or inhibitory effects. A disadvantage of this strategy
is the short half-life and high clearance rate of growth factors
in vivo
, especially when

Curr. Issues Mol. Biol. 2025,47, 287 13 of 28
administered systemically [
142
,
153
]. The incorporation of growth factors into scaffolds for
BTE has significant potential to improve therapeutic outcomes [2,146].
Runx2 is a key transcription factor involved in regulating MSCs’ differentiation into
osteoblasts. Runx2 increases the expression of osteoblast-specific genes and initiates miner-
alization [
119
,
154
]. In addition, the activation of Runx2 prevents MSCs from differentiating
into other lineages, such as adipocytes [
155
]. Kang et al. transduced the Runx2 gene into
human MSCs using a lentiviral vector and transplanted them into mice. As a result, supe-
rior bone healing was demonstrated compared to unmodified controls. The transplanted
cells migrated to the fracture site and differentiated into osteoblasts to form new bone [
119
].
Runx2-transduced MSCs showed enhanced osteogenic properties and elevated expression
of Runx2, ALP, OCN, and BSP
in vivo
in a model of cranial defect [
156
]. Runx2 may increase
the efficiency of bone healing, which may be clinically relevant, but its overexpression
has been associated with human malignancies in various cancers, indicating a need for
caution [119].
BMPs, the first identified factors inducing bone regeneration
in vivo
, are crucial pro-
teins for regulating the commitment of MSCs to osteoblastic fate [
136
]. BMPs are involved in
MSCs’ differentiation into chondrocytes and osteoblasts [
108
], as well as in that of osteopro-
genitors into osteoblasts. BMPs are produced by osteoprogenitor cells, MSCs, osteoblasts,
and chondrocytes and are mainly found in the extracellular matrix of bone. BMP signaling
can stimulate and play a role in almost every step of osteoblast differentiation and ultimate
cell maturation [
134
,
157
]. Among BMPs, BMP-2, 4, 5, 6, 7, and 9 are necessary for bone for-
mation, mainly by activating the Runx2, OCN, and OSX transcription factors [
116
,
119
,
158
],
and there is evidence of an increase in their efficiency when they are heterodimeric forms
of the BMP-4/-7 and BMP-2/-7 type,
in vitro
and
in vivo
[
116
,
119
]. Conversely, BMP-3
and BMP-13 are exceptions and exhibit an inhibitory function in osteoblast–osteogenic
differentiation [
159
,
160
]. BMP-2 is expressed on the first day of fracture healing to stimu-
late MSC differentiation, and BMP-6 and -9 are expressed at later stages in animal model
studies [
158
]. BMPs also play other roles in the healing process, such as stimulating the
synthesis and secretion of other bone-related and angiogenic factors, directly activating
endothelial cells for angiogenesis and regulating callus formation
[119,158]
. For example,
BMPs have been shown to induce new bone formation at the bone defect site [
142
], as the
process includes the initial inflammatory phase, soft callus formation, mineralization, and
bone remodeling [161].
In clinical trials, recombinant human BMP-2, BMP-4, BMP-6, BMP-7, and BMP-9 have
been reported to stimulate local bone regeneration by signaling MSCs’ differentiation into
osteoblasts [
162
,
163
]. BMP-2 and -7 have received special attention, as they are FDA (Food
and Drug Administration)-approved for bone regeneration applications [142,164]. BMP-2
is one of the most studied factors in MSCs regarding bone tissue regeneration and bioengi-
neering; even with its short expression, it strongly positively regulates OCN expression.
BMP-7, in turn, induces the expression of osteoblast differentiation markers, such as ALP,
as well as collagen-I synthesis and acceleration of mineralization [
165
]. Meanwhile, BMP-9
has been found to strongly induce the osteogenic differentiation of MSCs by activating
the canonical Wnt/
β
-catenin/Runx2/OCN axis, and Smad/JNK signaling induces the
osteogenic differentiation of mesenchymal C3H10T1/2 cells [
166
–
168
]. Interestingly, BMP-9
has a synergistic effect with TGF-
β
in the late stages of the osteogenesis of MSCs [
136
,
168
].
BMP2-iPSC-MSCs seeded on a scaffold for BTE showed almost twice higher mineral-
ization and ALP activity in vitro compared to the control [169].
The inclusion of BMP-4-containing nanoparticles in bioink composition had a dual
effect on a rat model of cranial defect: (1) it induced the secretion of BMP-2 by M2 type

Curr. Issues Mol. Biol. 2025,47, 287 14 of 28
macrophages; (2) BMP-4 and induced BMP-2 enhanced the osteogenic differentiation of
BM-MSCs and further accelerated bone repair [170].
TGF-
β
is a potent chemotactic stimulator, enhancing the proliferation of MSCs, pre-
osteoblasts, chondrocytes, and osteoblasts. TGF-
β
initiates signaling for BMP synthesis in
osteoclast cells, inhibits osteoclast activation, and stimulates osteoclast apoptosis. TGF-
β
and PDGF, released by activated platelets in the early stages of fracture healing, induce the
migration, activation, and proliferation of MSCs along with angiogenesis and the inflamma-
tory response. TGF-
β
signaling appears activated at early stages of osteogenesis, promoting
the acquisition of an immature osteoblastic phenotype by MSCs/osteoprogenitors while
inhibiting further osteoblast maturation, bone mineralization, and transition to osteo-
cytes [
134
,
136
,
171
]. TGF-
β
and BMP-2 are required for normal fracture healing. MSCs do
not differentiate into the osteogenic lineage without these factors, inhibiting healing [
172
].
However, the osteoinductive potential of TGF-
β
is limited and has shown various side
effects, thus limiting its clinical use for bone regeneration [119,158,172].
Histone demethylase JMJD3 positively regulates MSC differentiation into osteoblasts
in both intramembranous and endochondral bone formation. There is evidence that JMJD3
induces osteoblastic differentiation by stimulating transcription factors Runx2 and OSX
and controls the expression of genes related to bone formation [
119
,
172
]. Homozygous
deletion of JMJD3 strongly suppresses osteoblast differentiation and bone ossification in
mice [173].
IGF-1, released from the bone matrix, is involved in regulating MSCs’ differentiation
into osteoblasts [
174
] due to the activation of mTOR during the bone remodeling pro-
cess [
175
]. On the other hand, impaired IGF-1 signaling in MSCs decreases bone formation.
Knockout of IGF-1 in mouse MSCs impairs osteoblast differentiation and reduces trabecular
bone formation [174].
As for VEGFs, in the process of bone repair, angiogenesis precedes the onset of
osteogenesis. A combination of angiogenic VEGF, cell-recruiting PDGF, and osteogenic
BMPs has been investigated and demonstrated a synergistic effect that is more beneficial
for bone repair than either factor delivered alone [142,176].
Cyclin-dependent kinase 1 (CDK1) promotes the osteogenic differentiation of human
MSCs through the phosphorylation of enhancer of zeste 2 polycomb repressive complex 2
subunit (EZH2). Knockdown of CDK1 with three different shRNAs blocked osteogenic
differentiation and suppressed osteogenic markers, including Runx2 and OPN [177].
Regarding Stromal cell-derived factor 1 (SDF-1), it has been demonstrated that SDF-1
enhances the differentiation of MSCs into osteoblasts, which is mediated by the BMP
signaling pathway. Cells cultured in an osteoinductive medium with SDF-1 showed higher
ALP activity than cells cultured in an osteoinductive medium alone [
178
]. In addition, the
disruption of SDF-1 signaling decreased bone nodule mineralization and inhibited the
BMP-2-induced early expression of Runx2 and OSX, which are regulators of osteogenesis.
SDF-1 is upregulated at sites of injury, particularly the periosteum, and activates the
CXCR4 receptor on MSCs, promoting regeneration [
178
]. Blocking the SDF-1/CXCR4
axis or adding SDF-1 significantly affected BMP-2-induced ALP activity and osteocalcin
synthesis [179].
Transcriptional coactivator with PDZ-binding motif (TAZ) is a Runx2 transcriptional
coactivator for the osteocalcin gene in MSCs and suppresses PPAR
γ
-dependent gene
transcription. TAZ promotes MSC differentiation into osteoblasts and inhibits PRAR
γ
from facilitating MSCs’ differentiation into adipocytes [
180
]. TAZ co-activates Runx2 in
cells, which is critical for osteoblast differentiation. Experiments conducted by Hong et al.
showed that the differentiation of MSCs into osteoblasts depends on Runx2 and TAZ, and
their regulation may be an approach to bone formation [119].

Curr. Issues Mol. Biol. 2025,47, 287 15 of 28
Fucoidan, a sulphated polysaccharide extracted from brown seaweed, induces MSC
proliferation and promotes osteoblast differentiation through JNK- and ERK-dependent
BMP2-Smad 1/5/8 signaling in human MSCs. Fucoidan significantly increased ALP
activity and osteocalcin and BMP-2 levels associated with bone mineralization [
154
]. It
facilitates calcium accumulation and the regulation of osteoblast-specific genes, including
ALP, Runx2, alpha-1 collagen-I, and OCN. Fucoidan also induces BMP-2 expression and
stimulates the activation of extracellular signal-related kinase (ERK), c-Jun N-terminal
kinase (JNK), and Smad 1/5/8 by increasing phosphorylation. Smad signaling is known
to mediate the effects of BMPs, which are involved in bone formation signaling pathways.
The impact of fucoidan on osteogenic differentiation was inhibited by BMP-2 knockdown
and by ERK and JNK inhibitors, indicating that fucoidan affects differentiation through the
BMP2-Smad 1/5/8 signaling pathway by activating ERK and JNK [119,181].
IL-20 is involved in inhibiting osteoblast differentiation and maturation and increasing
osteoclast differentiation.
In vitro
, IL-20 increases sclerostin and decreases OSX, Runx2, and
osteoprotegerin (OPG), inhibiting osteoblast formation. IL-20 deficiency reduces fracture
healing time by limiting the inhibitory effects of IL-20 on osteoblastic differentiation from
MSCs and osteoprogenitor cells. IL-20 showed a significant correlation with sclerostin in
patients with bone fractures and osteoporosis. Sclerostin also inhibits the differentiation,
proliferation, and function of osteoblasts [
138
]. Anti-IL-20 monoclonal antibody 7E in
a mouse model increases bone formation at the fracture site, showing its potential as a
therapeutic agent for bone fractures [119,182].
The need for more understanding of MSCs’ therapeutic actions and their fate after
transplantation [
183
,
184
] has made the requirements for therapeutic pre-optimization,
such as optimal cell number, cell phenotype, maturity, and the mechanical properties of
tissue-engineered grafts, still challenging to define [107,185,186].
Currently, bioprinting research in orthopedics is focused on the osteogenic differ-
entiation of MSCs [
62
]. In 2017, Benning et al. demonstrated an enhanced proliferation
and osteogenic differentiation of drop-on-demand printed MSCs in collagen and fibrin
hydrogels containing hydroxyapatite [
187
]. Byambaa et al. developed cylindrical elements
made of GelMA hydrogels enriched with VEGF and loaded with BM-MSCs and HUVECs.
Bioprinted constructs have maintained cell viability and proliferation and successfully
induced osteogenic differentiation [
51
]. Aghajanpour et al. developed a self-oxygenating
bioprinted scaffold enriched with growth factors and loaded with BM-MSCs. The incor-
poration of both BMP-2 and calcium peroxide nanoparticles resulted in the upregulation
of Runx 2, Collagen type I alpha 1, and OCN genes compared to internal references in
osteogenic media [
188
]. In a study by Chai et al. [
189
], it was shown that scaffolds combin-
ing MSCs and osteoinductive factor BMP-2 had a better chance to regenerate bone defects
than scaffolds with cells/factor alone. These and some other examples are summarized in
Table 2.

Curr. Issues Mol. Biol. 2025,47, 287 16 of 28
Table 2. Some examples of bioprinted bone models with an added osteoinductive factor.
Cells Bioink Bioprinting Technique Growth Factor Study Type
(In Vitro/In Vivo) Bone Defect Summary of Findings REF
Rat BM-MSCs
GelMA/gelatine/PEG/0.4%
MSNs composite hydrogels
Extrusion-based 3D
bioprinting;
thermo-crosslinking;
photo-crosslinking of
GelMA
1µg/mL−1BMP-4 added
to 1 mg mL−1
MSNs-
Mesoporous silica
nanoparticles
Implanted subcutaneously
on the back of C57BL/6
mouse
Calvaria defect in diabetes
mellitus rats
GelMA/gelatine/PEG/MSNs composite
bioinks showed satisfactory printability,
mechanical stability, and biocompatibility.
The sustained release of BMP-4 from MSNs
induced M2-type macrophage polarization and
thereby inhibited inflammatory reactions.
Loading of BMP-4 and secretion of BMP-2 by
M2 type macrophages promoted the osteogenic
differentiation of BM-MSCs and further
accelerated bone repair in DM bone defects.
[170]
Human BM-MSCs Alginate hydrogel Extrusion-based 3D
bioprinting; ionic
crosslinking
Calcium peroxide
nanoparticles (CPO NPs),
BMP-2 (BMP2 NPs) In vitro -
The viability of encapsulated hBM-MSCs was
increased and osteogenic differentiation was
improved.
Applying a sustained-release formulation of
BMP-2 resulted in greater improvement in
hBM-MSCs’ osteogenic differentiation and in
the upregulation of RUNX2, OCN, and
COL1A1 genes.
[188]
Rat BM-MSCs GelMA hydrogel,
photo-crosslinked under
UV light (365 nm) for 30 s
Injected and
UV-crosslinked at the site of
injury BMP-2 SPF male SD rats Distal femur defect,
diameter 3 mm, depth of 2
mm
A photo-crosslinked BM-MSCs-BMP-2-GelMA
bioactive hydrogel scaffold effectively
promotes BM-MSC osteogenic differentiation
and bone tissue regeneration.
The active scaffold released about 70% of the
BMP-2 in the first week, which continuously
stimulated the adhesion and osteogenic
differentiation of BM-MSCs inside and outside
the scaffold.
[189]
Pre-osteoblast cell line
MC3T3-E1 GelMA/HAMA hydrogel
loaded with OGP Photo-crosslinking,
injectable Osteogenic growth peptide
(OGP) In vitro - The hydrogel promoted cell proliferation and
adhesion and increased osteogenic-related
gene and protein expression in vitro. [190]
Rat BM-MSCs GelMA hydrogels and
porous CaCO3
microspheres (CMs)
Injected into the site of bone
defect BMP-2 In vitro and
SD rats Skull defects
Rapid osteogenesis was induced, mainly
involving MSC recruitment and differentiation
in the later stage.
The inflammatory response was balanced;
macrophage polarization was modulated.
Appropriate and timely modulation of bone
healing process, such as the early inflammatory
stage and the later osteogenic stage, was crucial
to the healing of bone defects.
[191]
Rat BM-MSCs
MC3T3-E1 cells
HUVECs
Pearl powder (PP) hybrid
fish gelatine methacrylate
(GelMA)
Microfluidic-assisted 3D
printing technology VEGF Rats Skull defects
Controlled release of VEGF enables the scaffold
to promote angiogenesis.
A synergic effect of osteogenesis and
angiogenesis was seen.
[192]

Curr. Issues Mol. Biol. 2025,47, 287 17 of 28
Table 2. Cont.
Cells Bioink Bioprinting Technique Growth Factor Study Type
(In Vitro/In Vivo) Bone Defect Summary of Findings REF
human BM-MSCs
Polypyrrole-grafted gelatin
methacryloyl (GelMA-PPy)
with triple crosslinking
(thermo-photo-ioni-cally)
Extrusion-based 3D
printing Microcurrent stimulation
(250 mV/20 min/day). In vitro Printed full-thickness rat
bone model
Three-dimensional-bioprinted hBM-MSCs
highly expressed gene hallmarks for
NOTCH/mitogen-activated protein kinase
(MAPK)/SMAD signaling while
downregulating the Wnt/β-Catenin and
epigenetic signalling pathways during
osteogenic differentiation for up to 7 days.
[193]
BM-MSCs
Photo-crosslinked
biomimetic methacrylated
gelatin (Bio-GelMA)
hydrogel
GelMA-BM-MSC
suspension added into
PDMS mold and
photocured as scaffold
-In vitro and
rats Segmental bone defect
The BM-MSC-carrying GelMA hydrogel
scaffold has good mechanical properties and
biological compatibility.
It promotes the regeneration of bone and blood
vessels, improves the mechanical strength of
bone defects, and effectively promotes the
repair of bone defects.
[194]
iPSC-derived cells via
neural crest or mesoderm
overexpressing BMP-6
CellInk Bone or GelXA Bone
Cellink Bio X™ 3D
bioprinter (in vitro tests);
In vivo bone defect was
filled with spatula with
ink/cell mixture and
crosslinking agent was
added
BMP-6 NOD/SCID mice Cranial bone defect; frontal
and parietal bones
The combination of bioprintable bioink and
BMP-6 transfected iNCC-MPCs is capable of
stimulating bone regeneration. [195]
MSCs
Matrigel, fibrin, collagen,
gelatine, and
gelatine/alginate at various
hydrogel concentrations
Drop-on-demand (DoD)
printing HA In vitro -
The inclusion of HA enhanced the proliferation
and osteogenic differentiation of MSCs and
prevented the degradation of fibrin in vitro. [187]
MSCs
HUVEC
Naturally derived hydrogel
GelMA Extrusion-based
direct-writing bioprinting Silicate nanoplatelets and
VEGF In vitro Bone-like tissue constructs
containing a perfusable
vascular lumen
Encapsulated hMSCs formed a mature bone
niche after 21 days of culture under the
medium perfused condition. [51]

Curr. Issues Mol. Biol. 2025,47, 287 18 of 28
7. Clinical Challenges for Bioprinted MSC Applications
The idea of BTE and 3D printing applications in the clinic dates back to the late 1990s
and was initially utilized for printing dental implants, customized prostheses, and kidney
bladders [
88
], but despite the fact that almost 40 years have passed, BTE has still not
reached its dreamed aim, since only cell-free 3D printing strategies for bone repair have
reached clinical application [196]. Nowadays, cell-free 3D printing is used in surgery, and
orthopedic applications account for almost half of its total uses [
34
,
197
]. Three-dimensional
printing technology has a wide range of development prospects in orthopedics, but other
disciplines like otosurgery and dental and maxillofacial surgery are also in demand for
bioprinted implants for regenerating patients’ bone defects.
Stem cell therapy presents a lot of potential for regenerative procedures in dentistry,
utilizing the ability of stem cells to differentiate into different types necessary for tissue
repair [
198
]. MSCs, particularly those derived from dental tissues such as the dental pulp,
apical papilla, and periodontal ligament, are of particular interest due to their accessibility
and osteogenic potential [
198
–
200
]. Induced pluripotent stem cells (iPSCs) also show
significant potential [
201
]. These cells have the capability to regenerate various oral tissues,
including bone, dentin, and periodontal structures, offering solutions for conditions such
as tooth loss and periodontal disease [202].
Over time, 3D bioprinting has established itself as a powerful tool for creating cus-
tomized scaffolds that mimic the complex architecture of dental tissues. Systematic reviews
on the application of stem cells in maxillofacial regeneration have shown promising re-
sults, particularly with MSCs, though they emphasize the need for standardized protocols
and long-term data [
201
]. Another systematic review explores the current clinical appli-
cations of stem cell therapies in facial reconstruction and regenerative surgery [
203
]. A
novel three-dimensional construction strategy for generating mineralized bone structures
using BM-MSCs is presented in [
47
].
In vitro
and
in vivo
studies on bioprinted scaffold
models demonstrate their potential for regenerating bone, periodontal tissue, dentin, and
pulp [
204
]. These models often incorporate growth factors and other bioactive molecules
that enhance cell differentiation and tissue formation.
The combination of dental pulp stem cell-derived exosomes and xenografts has proved
to be a promising strategy for enhancing new bone formation and regenerative scores in
repairing critical-size defects [205].
Stem cell therapies and 3D bioprinting are transforming the field of dentistry and offer-
ing new opportunities for the regeneration of damaged or lost tissues. Despite significant
advancements, further research is required to optimize bioprinting techniques, develop
the most effective bioinks, standardize stem cell protocols, and validate long-term clinical
outcomes. This approach has the potential to revolutionize dental treatments and improve
patient care.
8. Conclusions and Perspectives
Bone defects can deprive a person of the most basic support, which can lead to a range
of problems. The treatment of bone defects remains a clinical challenge, and autologous
transplantation is limited and cannot cover the needs. With the continuous research on
bone tissue engineering, 3D-printed scaffolds have emerged; they solve the shortage of
grafts and offer the option to be personally designed. Their osteogenic effect still cannot
meet the needs of clinical treatment. Cells must proliferate to establish cell–cell connections
and communicate with each other, as well as secrete extracellular matrix components
and perform specific biological functions to integrate into the host tissue effectively [
88
].
Therefore, 3D-printed scaffolds loaded with growth factors were born [34].

Curr. Issues Mol. Biol. 2025,47, 287 19 of 28
Current technical capabilities are aimed at the independent or combined application
of three new approaches: the construction of an “artificial intercellular substance” with
appropriate mechanical parameters and an individualized shape (3D printing) which will
overcome the shortcomings of transplantation and facilitate surgical activity; the application
of different cell types, mainly MSCs, with the idea of the subsequent biodegradation
of the implant and the replacement of the implant with the patient’s own tissues; and
the acceleration and management of regeneration by the addition of biologically active
substances. Building a structure with 3D printing is associated with many requirements
for the material. In
in vitro
experiments, substances, whether biodegradable or not, are
selected when they are suitable for the various 3D printing methods and meet certain
biophysical parameters (strength, elasticity, weight, etc.). When cells are embedded into
the structure during the “printing” process (so-called bioprinting), the requirements for the
material become very limited, and currently, there are a relatively limited number of bioinks
for “bone tissue”, mainly with collagen/gelatine [
41
] and cellulose microfilaments [
206
]. A
limiting factor is also the desire for the structure to include biologically active substances,
which are relatively unstable under the conditions of “printing”.
Despite intensive research, only cell-free prints are used in clinical practice as of now.
In quite a lot of studies, a cell-free scaffold is printed first, and later, in a post-printing
modification, it is loaded with cells. An alternative is to mix the bioink with cells, sometimes
with factors as well, but this composite is injected into the site of the defect in the laboratory
model, and it is not bioprinted. In the literature, plenty of osteogenic factors have been
studied, some of them listed and briefly described in this review. A very limited number of
them (BMP2, BMP4, BMP6, and VEGF) have already been incorporated in BTE scaffolds.
Future research on osteoinductive factors may lead to the development of novel bioinks that
promote bone regeneration more effectively. Future extensive research on truly bioprinted
bioink/cell/factor composites is needed before their routine clinical application for bone
defect repair.
Author Contributions: Conceptualization, M.K. and M.M.; writing—original draft preparation, M.K.;
writing—BTE in clinical practice, R.S., M.R. and S.L.; writing—review and editing, M.K., S.L. and
M.M.; visualization, M.K.; project administration, M.M. All authors have read and agreed to the
published version of the manuscript.
Funding: This research was funded by FNI-Bulgaria, grant number KP-06-H53/7; M.M. and S.L. are
participants in InfraACT, grant agreement DO1-178/29.7.2022, National Roadmap for RI.
Conflicts of Interest: The authors declare no conflicts of interest.
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