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Cells 2024, 13, 1638. https://doi.org/10.3390/cells13191638 www.mdpi.com/journal/cells
Review
Bioprinting of Cells, Organoids and Organs-on-a-Chip
Together with Hydrogels Improves Structural and
Mechanical Cues
Claudia Tanja Mierke
Faculty of Physics and Earth System Science, Peter Debye Institute of Soft Matter Physics, Biological Physics
Division, Leipzig University, 04103 Leipzig, Germany; claudia.mierke@uni-leipzig.de
Abstract: The 3D bioprinting technique has made enormous progress in tissue engineering, regen-
erative medicine and research into diseases such as cancer. Apart from individual cells, a collection
of cells, such as organoids, can be printed in combination with various hydrogels. It can be hypoth-
esized that 3D bioprinting will even become a promising tool for mechanobiological analyses of
cells, organoids and their matrix environments in highly defined and precisely structured 3D envi-
ronments, in which the mechanical properties of the cell environment can be individually adjusted.
Mechanical obstacles or bead markers can be integrated into bioprinted samples to analyze mechan-
ical deformations and forces within these bioprinted constructs, such as 3D organoids, and to per-
form biophysical analysis in complex 3D systems, which are still not standard techniques. The re-
view highlights the advances of 3D and 4D printing technologies in integrating mechanobiological
cues so that the next step will be a detailed analysis of key future biophysical research directions in
organoid generation for the development of disease model systems, tissue regeneration and drug
testing from a biophysical perspective. Finally, the review highlights the combination of bioprinted
hydrogels, such as pure natural or synthetic hydrogels and mixtures, with organoids, organoid–cell
co-cultures, organ-on-a-chip systems and organoid-organ-on-a chip combinations and introduces
the use of assembloids to determine the mutual interactions of different cell types and cell–matrix
interferences in specific biological and mechanical environments.
Keywords: collagen; viscosity; cell–matrix bidirectional interaction; 4D bioprinting; organoids;
cancer; polymers; stiffness; assembloids; tumoroids
1. Introduction
In medicine, biomedical and biophysical research, it is commonly known that there
is and will be a continued massive demand for tissues and organs for transplantation,
experimental drug screening and the fundamental analysis of developmental and dis-
eased processes of cells or cell clusters in tissues. As there are not enough organs available
for patients, there is no question that tissues and organ models must be developed and
engineered for research, which is the overall purpose for organoid cultures in general. For
many decades, biologists have employed two-dimensional (2D) culturing, veterinary
models or dead bodies to gain precious knowledge about disease mechanisms, drug
screening and safety (toxicity) studies, but the transfer to humans is debatable with regard
to its accuracy, validity, significance and repeatability [1,2]. In addition, due to their lim-
ited efficacy, more durable, perhaps more radical, replacement methods are needed for
existing cell therapy approaches for the treatment of chronic diseases. An extremely im-
portant task for tissue engineering is to replace or possibly avoid animal testing alto-
gether, which is also a commonly defined well-known aim in the field. This is highly de-
sirable from a bioethical point of view and can also be the answer to practical problems
Citation: Mierke, C.T. Bioprinting of
Cells, Organoids and Organs-on-a-
Chip Together with Hydrogels
Improves Structural and Mechanical
Cues. Cells 2024, 13, 1638.
https://doi.org/10.3390/cells13191638
Academic Editor: Alexander E.
Kalyuzhny
Received: 20 August 2024
Revised: 25 September 2024
Accepted: 1 October 2024
Published: 1 October 2024
Copyright: © 2024 by the author. Li-
censee MDPI, Basel, Switzerland.
This article is an open access article
distributed under the terms and con-
ditions of the Creative Commons At-
tribution (CC BY) license (https://cre-
ativecommons.org/licenses/by/4.0/).
Cells 2024, 13, 1638 2 of 87
arising from species-specific variations in cell function and tissue organization [3]. More-
over, realistic three-dimensional (3D) tissue models are increasingly needed for toxicolog-
ical in vitro screening and drug development [4]. Advanced tissue engineering via bi-
oprinting of 3D organoids is occupying an emerging key role in biomedical, cell biological
and biophysical or mechanobiological research [4,5].
Of the tissue engineering techniques, 3D bioprinting is set to revolutionize the bi-
omanufacturing of tissues and organs [6]. Scaffolds have been utilized as a possible vehi-
cle for producing artificial organs and tissues prior to the emergence of the 3D printing of
living cells [7]. While scaffolds are not a novel tissue engineering methodology, their tra-
ditional techniques and structures entail multiple constraints for the generation of effica-
cious tissues and organs. The main restriction of scaffolds is that they are not able to imi-
tate the inherent functionalities of an extracellular matrix (ECM), as the precise mecha-
nisms of their functioning are still not completely comprehended. In addition, traditional
methods of scaffold production also fail to replicate the natural mechanical characteristics
that are essential for correct biological performance [7], such as the issue of inhomogeneity
of natural ECMs, which can only be partially reproduced in conventional hydrogel cell
culture models [8]. In contrast, the 3D bioprinting process for tackling these issues is built
on the principles of biomimicry, autonomous self-organization and microtissues, which
are typically intricate in vivo tissues composed of several simpler units where the compo-
site structure and function form the complete picture [9,10].
The combination of in vitro cultured 3D cell structures such as organoids, which can
self-organize and mimic real organs in structure and function, with bioprinting techniques
appears to be highly suitable for rapidly advancing model systems and offers enormous
potential for physiological and pathological questions. Organoids come from stem cells
or tumor tissue (tumor organoids) taken from patients and are grown in vitro in a specific
3D microenvironment. The integration of tumor organoids into 3D-printed tumor models
is currently in its fledgling stages but provides new opportunities for more precise tumor
microenvironment (TME) reconstitution. The TME is composed of non-malignant cells
(immune cells, stromal cells and (tumor) endothelial cells) and ECM protein contributing
to a 3D ECM scaffold [11]. Non-malignant cells can interfere among each other and with
cancer cells, such as those of solid tumors, which consequently critically affect cancer bi-
ological analyses [12]. In addition, the TME is intricate and required for tumor growth,
and hence, simplifying tasks are needed for gaining insights [13]. The cutting-edge tech-
nique of 3D bioprinting accurately mimics collective cell performance, individual patient-
specific physiological properties, and the accurate monitoring of geometric, biophysical,
and biochemical parameters in the TME, such as cell-derived products released via extra-
cellular vesicles. Thereby, the mechanical environment of cells and 3D organoids can be
rebuilt [14]. Moreover, by continuing to implement organ-on-a-chip technology in bio-
medical and biophysical research, fluid dynamics and immune cells, such as macrophages
and neutrophiles [15], can be incorporated in an easy manner, which can then remodel
the external environment of cells by mechanical signals, and thus, for example, enhance
the accuracy of drug screenings.
Tumor organoids are increasingly deployed as in vitro models in cancer research due
to their ability to recapitulate the complexity of tumors [11]. Such patient-derived cancer
cell assemblies can closely replicate key tumor characteristics, comprising tumor-like cell–
cell junctions, heterogeneous cell populations, (epi-)genetic environment and parental tu-
mor growth characteristics [16]. Following the triumph of the first tumor organoids de-
rived from colorectal cancer tissue by Sato et al. [17], organoids of diverse cancer types
have been evolved with considerable impact for the modeling of cancer heterogeneity and
personalized medicine. As organoid culturing systems have progressed, several trials
have established biobanks of patient-derived tumor organoids (PDOs) for bladder [18],
brain [19], breast cancer [20], cervical cancer [21], colorectal cancer [22], esophageal cancer
[23], gallbladder cancer [24], gastrointestinal cancer [25], glioblastoma (GBM) [26], head
and neck cancer [27], liver cancer [28], lung cancer [29], ovarian cancer [30], pancreatic
Cells 2024, 13, 1638 3 of 87
cancer [31], prostate cancer [32], thyroid cancer [33] and multiple other cancer types. PDO
biobanks offer a resilient foundation for patient-specific high-throughput screening of
therapies, incorporating chemotherapy, immunotherapy and radiotherapy [20,29,34], and
warrant multiple trials on tumor advancement and TME hallmarks [28,35].
Besides heterogenous cancer cells, the value of including heterogeneous stromal cell
types in in vitro tumor models is increasingly acknowledged [11,36], and it has been
shown to impact mechanical cues of the TME [36]. In turn, tumors can reprogram nearby
or remote (via extracellular vesicle release) supporter cells into activated subtypes to sus-
tain cancer progression, promote resistance to chemotherapy and circumvent the immune
defense reaction [37]. Integrating these heterogeneous stromal cell types and regulation
of their switch aids in obtaining substantial tumor-specific expression data profiles in in
vitro tumor organoid models. Whereas standard culture conditions of PDOs eliminate
specific key features of cancers, innovative co-cultures of PDOs with matching cancer-
associated fibroblasts (CAFs) sourced from the same patients have been shown to increase
transcriptome stringency and the direct subtype-specific expression of stromal genes [38].
Thereby, the mechanical cues of the environment can be mirrored, as CAFs can stiffen the
surrounding ECM via lysyl oxidase (LOX) enzyme crosslinking of collagen I [39] and
hence, indirectly alter cellular shape, mechano-phenotype and function [40]. Moreover,
the composition of the ECM and elevated ECM stiffness, which are both major contribu-
tors of cancer’s progressive and invasive capacity, whereby the TME stiffness as emerged
as a biomarker for cancer [41]. Consequently, the ECM stiffness impacts the outcome of
cancer therapies [42]. Similarly, a co-culture model that paired PDOs with endogenous
tumor-infiltrating lymphocytes as cohesive entities has been found to adequately capture
the tumor-infiltrating immune milieu [43]. These organoid frameworks also fostered the
emergence of new immunotherapies and made it possible to examine the tumor reaction
to anti-programmed cell death protein 1 (PD-1) and anti-programmed death ligand 1 (PD-
L1) therapies through uncoupling the cancer-infiltrating and cancer-surrounding ele-
ments, such as tumor stroma modifications or mechanical characteristics alterations [44].
Tumor organoids hold tremendous power for rebuilding 3D architecture and heter-
ogeneous cellular elements. They offer distinct advantages over other culture models,
among them an improved capacity to mimic the physiological and pathological condition
of tumor organs, a moderate expense and an improved compatibility with several new
emerging technologies. However, they are inherently restricted in their capacity to mimic
other crucial elements of the TME, like tumor-specific biochemical/biophysical character-
istics, anatomical sizing, hierarchical blood/lymphatic vasculature and fluid dynamics
[45,46]. The application of 3D bioprinting with tumor organoids can transcend these lim-
itations to create comprehensive, concise models with greater clinical efficacy [1,47,48].
Moreover, this technique can reduce the animal testing for drug screening approaches in
the future [1]. Alongside the applications of 3D printing in cancer research, tissue engi-
neering approaches and the generation of functional biological structures that are able to
substitute or rebuild compromised tissue have continued to advance in recent years and
have affected a broad spectrum of medical disciplines, encompassing the bioprinting of
vascular channels [49], osseous implants [50], dermal transplants [51], intestinal trans-
plants [52] and heart tissue [53].
The purpose of the review is to discuss the recent progress in 3D bioprinting and new
4D bioprinting approaches through which the structural and mechanical cues of ECM en-
vironments can be rebuilt in a more precise and reliable manner from a biophysical or
mechanobiological perspective. These techniques also enable the dynamic restructuring
of mechanical cues. Bioprinting technology overcomes certain constraints of organoid
production, allowing it to be increasingly utilized in drug testing, regenerative medicine,
cancer research, and mechanobiology. The focus of the overview is on its use and appli-
cations in cancer research; however, other biological applications in healthy and diseased
individuals are also included to illustrate the broad application of this technique. Special
attention is paid to the interaction of 3D-printed cells and the surrounding hydrogels in
Cells 2024, 13, 1638 4 of 87
3D cell printing to produce organoids. Aspects covered in the review comprise a brief
overview of bioprinting techniques, including extrusion bioprinting, microextrusion,
inkjet bioprinting and stereolithography (SLA), laser-induced forward transfer (LIFT), an
introduction to the role of bioinks, the physiochemical and biological properties of poly-
meric hydrogels (natural, synthetic and hybrid), scaffold-free and scaffold-based bioprint-
ing, the combination of organoids or organ-on-a-chip and bioprinting techniques, the role
of cell alignment in printed scaffolds and finally a brief outlook on future developments
in the field of mechanobiology. Subsequently, the relevance of organoids compared to the
organ-on-a-chip technology and important architectural factors for the design of integra-
tive organotypic tumor models are emphasized. In the future, the further development of
microfluidic systems, controlled mechanical stimulation, advanced organoid models and
four-dimensional bioprinting technology could help to create better bioprinted organoids.
2. Brief Overview of Major Bioprinting Techniques for Cells, Hydrogels
and Organoids
In this section, a brief and basic introduction to the complex field of major bioprinting
techniques is provided from a cellular biophysical perspective. The advanced reader is
referred to more detailed review articles [54,55]. The first fast layer-by-layer prototyping
technologies for the production of 3D designs was invented in the 1980s and was used
for SLA-based 3D printing [56,57]. New techniques such as selective laser sintering (SLS),
fused deposition modeling (FDM), laminate object manufacturing (LOM) and electron
beam melting (EBM) have all been pioneered for a broad array of materials [58], such as
metals [59], ceramics [60] and diverse thermoplastics [61]. The layer-by-layer fabrication
ability of intricate constructs is specifically beneficial for the creation of in vitro tissue-
engineered models consisting of cells and other biological materials. Different types of 3D
printing have been explored for the production of complex tissues like bone, cartilage,
heart, muscle and liver [57,62–65]. Particularly for tissue engineering purposes, there are
continuous novelties in the development of different synthetic and natural polymers
[31,66,67], nanomaterials [68–70], high internal phase emulsions (HIPEs) [71], ceramic
composites [72–74], decellularized extracellular matrix (dECM) materials [75,76] and con-
ductive materials [77–79] for 3D printing applications. Cells and bioactive molecules are
embedded in the bioink, which increases the complexity of the material and the overall
approach. It is easy to incorporate vascularization and compartmentalization into 3D or-
ganoids in a pre-designed manner, which requires changes to the printing process, such
as temperature, pH, printing speed and mechanical pressure when bioprinting cells [80].
Multiple 3D printing techniques have been advanced for the bioprinting process,
among them inkjet, (micro)extrusion, laser and SLA printing [81]. The focus lies on micro-
extrusion bioprinting processes, as these are most often utilized for hydrogel materials.
Microextrusion processes employ uninterrupted pneumatic pressure or mechanical forces
driven by motors or screws to eject the bioink out of the printing nozzle as a continuous
filament [55]. In contrast to 3D printing with a thermoplastic material, the printing param-
eters such as temperature, pH value and pressure as well as the material characteristics of
the bioink are severely limited when living cells are deployed for bioprinting [82].
These bioink materials often suffered from a general lack of biological performance,
problems with cell survival and the creation of intricate architectures. Although much
progress has been made in these initial pioneering efforts, a remaining obstacle in this
field is the suitability of bioinks to support cells at all stages of the printing process. In
particular, firstly, cell survival during suspension in the syringe; secondly, cell survival
during extrusion from the nozzle; thirdly, cell survival at all stages of material bonding
and fourthly, not least, survival as the finished construct ages and takes on a tissue-like
form after printing [83].
Cells 2024, 13, 1638 5 of 87
2.1. Microextrusion or Extrusion Bioprinting (Screw, Piston, Pneumatic)
Extrusion (or microextrusion) bioprinting technology is an upcoming technology in
which biomaterials are accurately layered with living cells (termed bioink) to form 3D
functional structures for tissue engineering (Figure 1).
Figure 1. Schematic sketch of a pneumatic extrusion bioprinting device.
Printability, the ability to build and maintain a replicable 3D structure, and cell via-
bility (surviving cells in the printing process) are two of the critical features of the extru-
sion bioprinting process [84]. Extrusion bioprinting is extensively utilized to produce cell-
integrated designs for tissue engineering with manufacturability and cell viability being
two key aspects [85]. The discrepancy between printed and engineered structures is a ma-
jor difficulty, and limits progress in mimicking native tissue organs or TMEs for use in
tissue engineering and cancer research [80]. Among the many factors that can influence
the printability of structures are the bioink characteristics, the settings of the printing tech-
nique and the shape of the structure [86]. An advantage of the extrusion bioprinting tech-
nique are the viscosity regulation of the bioink can improve the printing process [85]. In
addition, nanoparticles can be included for mechanical analysis [87], various types of
crosslinking, such as reversible, chemical, physical and enzymatic, multiple available bi-
oinks and the on-going development of new bioinks, printing of dispersed cells, cell sphe-
roids and tissue strands and it is widely used technique [88]. Major disadvantages of the
extrusion bioprinting technique are its limited resolution: cells cannot be precisely pat-
tered and organized. In addition, the bioprinting process could induce quantifiable cell
death caused by alterations in dispensing pressure, nozzle geometry, printing time and
bioink concentration. Moreover, the bioinks for extrusion bioprinting need to perform
“liquid to solid” transition at the right time. However, the application of cell spheroids is
limited, as they should not be too large, as otherwise the core cells will become inactive or
necrotic due to a lack of oxygen [89]. It is challenging for extrusion bioprinting to recreate
the blood supply reticulation. In addition, the extrusion of the bioink from the nozzle us-
ing pneumatic pressure or mechanical force by means of a piston or screw can mechani-
cally stimulate the cells and change their function [90]. The key benefit of extrusion bi-
oprinting over other processes is the capability to embed cells into the biomaterials for
printing structures, whereas the process-related forces may compromise the embedded
cells (or the viability of the cells)—another non-negligible concern in the field of bioprint-
ing [91]. Strain stress and shear stress represent two important process-related forces that
cause cell injury. Some key factors like needle type and size, the concentration of bioink
Pressure/ Force
Biomaterial ink,
such as collagen or
MatrigelTM and or
cells/ organoids
Extrusion
nozzle
Printed
structure
Cells 2024, 13, 1638 6 of 87
and dosing pressure contribute to cell deterioration. Despite numerous promising inves-
tigations of the printability and viability of cells, this research area is in its infancy and the
accurate identification of effective factors continues to be a fundamental concern for future
advances. A trial-and-error determination or improvement approach is costly, challeng-
ing, tedious and, at times, infeasible; thus, computer-aided techniques are emerging as
powerful instruments. Many interdependent factors are relevant to the improvement of
the bioprinting procedure [61,92]. Machine learning is a new technology that can be used
in the field of 3D bioprinting to greatly advance this technology [93]. The biggest challenge
in the further development of machine learning in the field of bioprinting is presently the
scarcity of existing data. For this reason, it is necessary to create a global database sharing
system for bioprinting. Sharing data due to the different brands of bioprinters and soft-
ware across the globe can raise many new questions [94]. Therefore, standardized data for
every bioprinter using similar open-source software for all the printers seem auspicious.
Machine learning, although new in the field of bioprinting, is anticipated to transform
bioprinting and thus tissue engineering in the coming years [93]. In addition, (micro)ex-
trusion bioprinting facilitates the production of heterogeneous structures exhibiting high
form accuracy by depositing a bioink possessing the targeted physicochemical and bio-
logical properties [95]. A novel semi-synthetic hydrogel, composed of gelatin methacry-
late and Pluronic F127, has been custom formulated to meet the demands of the (micro)ex-
trusion bioprinting technique [96]. The combination of the thermosensitive properties of
Pluronic with the crosslinking properties of gelatin methacrylate provides the compound
with a printability range offering good dimensional stability and chemical stability after
photocrosslinking [97], as revealed by a rigorous evaluation of printability using predic-
tive empirical models. The mechanical characteristics of the structures are similar to that
in soft tissue, which expands the scope of soft tissue engineering. The bioink has been
effectively used to produce multilayer porous assemblies that retain a high degree of cell
viability [95]. Interestingly, the spatial layout of the cells exhibited a high level of align-
ment following the direction of deposition [98]. Finally, this manufacturing process could
provide a versatile approach for the creation of 3D models with a predefined cellular ori-
entation [99]. In the specific case of tissue engineering, (micro)extrusion bioprinting has
been applied in various areas, from the production of vascular prostheses to skin grafts
toward 3D organoids with a vasculature [100,101]. This technique is based on the ejection
of ink through mechanical or pneumatic forces. Compared to other techniques, such as
inkjet, the advantage of this technique is that high-viscosity liquids and structures with a
very high cell density can be processed. (Micro)extrusion bioprinting involves extruding
the bioink straight in its gel phase, with no support structures or cross-linking substances
introduced to the material as it flows out [63]. Therefore, the printing approach is highly
reliant on the rheological characteristics of the bioink, which is made of a biomaterial hy-
drogel harboring living cells [102–105]. In this setting, a key difficulty is the design of ap-
propriate inks that exhibit both the desired properties of extrudability and stability after
printing [106]. Hence, multiple mixtures of natural and synthetic biomaterials with com-
plementary characteristics, such as shear thinning or strain hardening, have been ana-
lyzed to generate appropriate chemical, mechanical and biological attributes for intended
purposes [107,108]. Multicomponent inks have proven to be an ideal choice to tackle the
disadvantages of single-material formulations, like the restricted printing ability of natu-
ral polymers or the deficiency of cell-specific activity associated with synthetic polymers,
thereby incorporating the benefits of every ingredient and extending the range of biofab-
rication [109]. When various hydrogels are blended together, the final polymer blend can
benefit from different crosslinking mechanisms, including both chemical and physical
crosslinking modes, resulting in robust gel stability [57,110]. This approach has been
widely pursued and material systems with improved printability have been identified for
the manufacture of functional tissues featuring intricate architectures [107,111–115]. Plu-
ronic F127 (PF127), a poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide)
triblock copolymer, has been utilized as a sacrificial template medium in a multi-
Cells 2024, 13, 1638 7 of 87
component proprietary design to define the entire thermal performance of the bioink and
facilitate the straight laydown of low viscosity alginate [116,117], that is realized to have
inadequate extrudability in its non-crosslinked form [96], increasing the bulk porosity of
the hydrogel scaffold upon its withdrawal.
A new type of proprietary formulation of Pluronic F127 and gelatin methacrylate
(GelMA) has been presented that optimizes the printability and biocompatibility. GelMA
is a natural-based hydrogel of great potential because it is photohardenable by gelatin
functionalization with methacrylate groups and can be manipulated easily [97]. For in-
stance, its mechanical characteristics can be easily controlled by adjusting several key fac-
tors like the level of substitution and the amount of polymer, while its rheological charac-
teristics are mainly governed via the processing temperature and dosage of UV radiation
[118–120]. GelMA can also be an optimal biomaterial for tissue engineering purposes, as
it strongly mimics the tissue microenvironment because of the existence of natural cell
adhesion moieties and provides the expected amount of bioactivity anticipated from a
tissue engineering framework [121–124]. In (micro)extrusion-based bioprinting, the vis-
cosity limitations of the bioink, the gelling time and operating parameters, including pres-
sure and printing velocity, are critical for the deposition of fibers with a predefined shape
and scalable dimensions [125–127]. Besides other characteristics, an optimal bioink needs
to have a distinct shear thinning characteristic and be able to regain its viscosity at a stand-
still directly following extrusion [128].
2.2. Inkjet Bioprinting (Piezoelectric Actuator, Heater, Jet-Based, Drop-on-Demand)
The advantages of low price, high efficiency and high accuracy have ensured that
inkjet printing is widely used [129]. The inkjet bioprinting process involves the production
of small, size-adjustable droplets of liquid ink, which are applied to the surface of the
substrate at specific points (Figure 2).
Figure 2. Schematic sketch of a inkjet bioprinting device.
Besides its exceptional efficiency and high accuracy, inkjet technology possesses sev-
eral unique properties, such as drop-on-demand technology and non-contact material
feed [130]. As a result, it has also been utilized for printing electronics equipment compo-
nents, structural polymers, sol-gel materials, biomolecules and living cells since approxi-
mately the year 2000 [131–134]. After the concept of 3D printing was unveiled, inkjet print-
ing has been progressively customized for 3D manufacturing. Three-dimensional inkjet
printing technology has attracted a great deal of interest in numerous research fields, such
as polymer molding [61,135], nanocomposites [136], drug delivery systems [137], organ
and tissue engineering [55,138–140] and the generation of cancer model systems [141].
Inkjet printing is a pioneering 3D printing platform for cell printing with the purpose of
manufacturing whole tissues or organs. Based on this advancement, the new concept of
“inkjet bioprinting” is a branch of “bioprinting” that has been formerly characterized and
Thermal,
electromagnetic,
piezo or acoustic-
based printhead
Printed droplets
containing cells
and bioink
Coated platform
Cells 2024, 13, 1638 8 of 87
authorized [63,142,143]. In short, “inkjet bioprinting” refers to the manipulation, structur-
ing or assembling of biologically important substances, like generic biomaterials, biomol-
ecules and cells, in a prefabricated way to serve specific biological purposes using inkjet
fabrication technology. In addition, inkjet bioprinting has been integrated with these es-
tablished key technologies to promote their further development. Inkjet technology can
simultaneously deposit a wide spectrum of materials and cells to specific locations in a
non-contact, tailored way, providing the opportunity to construct intricate heterogeneous
biomimetic patterns. The droplet size generated by inkjet printing can reach the order of
picoliters, which is suitable for the high-precision positioning of micro-scale biological
entities in a digitized design. In addition, the diameter of the nozzle of inkjet printers is
generally around 50 μm, which falls in the same size range as cells and thus opens up the
possibility of printing cells and individual cells [144].
In 1988, collagen and fibronectin were printed with an thermal inkjet device [145]. In
1997, poly(lactic-co-glycolic acid) (PLGA) and poly-lactide (PLLA) scaffolds were fabri-
cated with 3D inkjet printing technique [146]. The frameworks were colonized with pri-
mary hepatic cells and endothelial cells, whereby the cell mixture attached to the frame-
work and a specific tissue pattern could be rebuilt. In 2000, denatured DNA was loaded
into the ink cartridge of an thermal inkjet printer and printed out as a text design [147].
The prepared sample was then hybridized with the printed DNA array, which led to the
formation of a distinctly visible hybridized array. This confirms the suitability of inkjet
printing for producing DNA arrays and illustrates the huge, tremendous possibilities of
inkjet printing [148]. Bioactive proteins can also be applied by inkjet printing and their
folding conformation can be preserved post-printing [149]. A protein-containing buffer
suspension has been deposited on a glass plate to form a protein microarray with spots of
approximately 150–200 μm in diameter. This technique has been employed to investigate
protein–protein, enzyme–substrate and protein–small molecule interactions, demonstrat-
ing both the activity of the printed proteins and the wide range of potential applications.
In 2003, Biotin was printed with an inkjet printer and the idea of printing cells with inkjet
technology was raised for the first time [150]. The first printing process with living organ-
isms was performed with bacteria [151]. The bacterium Escherichia coli was selected to
print in precise designs and density gradient arrangements, demonstrating that cell ink,
like ordinary ink, can create specific designs. Mammalian cells are bigger than bacteria
and are easier to injure when being printed, so printing them is a greater difficulty. A
thermal inkjet printer was employed to print Chinese hamster ovary cells on soy agar gel
medium and collagen gel medium, reaching a cell survival rate of more than 90% [152].
In 2009, a half-heart with connected ventricles was printed [153]. Later, chondrocytes were
printed in an effort to create artificial cartilage [154]. Since natural animal tissues consist
of several types of cells with intricate architectures that cooperate to fulfill specific biolog-
ical functionalities, inkjet printing must be able to apply multiple cell types and materials
at the same time. In 2013, 3D inkjet printing of heterogeneous cells was implemented with
success [155]. Different cell types were transferred into different ink reservoirs of an inkjet
printer and then deposited layer by layer at specific locations to form intricate multicellu-
lar 3D hybrid structures. Subsequent in vitro and in vivo assays revealed that each cell
type remained viable and could perform physiological activities [156].
Although inkjet printing offers unique advantages for clinical use, there are several
limitations that hinder the further development of this technology [130]. Currently, only
a limited number of materials are suitable for inkjet bioprinting. A general disadvantage
of inkjet printing lies in the fact that the ink must be fluid to prevent clogging and that
droplets are created in the process. In addition, the requirements and specifications for
bioinks are much stricter than those for conventional inks, since it is not only biocompat-
ibility, degradability, mechanical characteristics and commercial advantages but also flu-
idity and viscosity that must be taken into account when selecting a bioink. Due to the
reduced viscosity prior to crosslinking and the auxiliary function post-crosslinking, hy-
drogel-based substances like alginate and calcium chloride, as well as acrylated
Cells 2024, 13, 1638 9 of 87
polyethylene glycol (PEG), are currently used extensively in 3D inkjet technology. Never-
theless, certain crosslinkers are toxic and are not approved for use in printing of cells.
Conversely, the bioink utilized in inkjet printing should have a fairly low concentration
to prevent blockages, which can render the printed 3D patterns unsuitable in terms of
geometry or performance [157]. The limited viscosity and restricted materials are there-
fore the key challenges of inkjet printing in the field of biotechnology [158].
The technique of inkjet bioprinting demands a reasonably low viscosity of the printed
composite and sufficient mechanical strength to retain its structural characteristics post-
printing. The existing photocrosslinkable media offer new possibilities. A cell-loaded bi-
oink comprising acrylated peptides and acrylated PEG was subjected to simultaneous
photopolymerization upon release [159]. The cell exhibited excellent viability and block-
ages were reduced to a minimum because of the reduced viscosity of acrylated PEG [160].
In addition, excellent mechanical characteristics were obtained, and the grafted human
mesenchymal stem cells remained in place in the printed pattern and formed homogene-
ous skeletal bone and cartilage. The strategy of inkjet printing with photocrosslinkable
inks employing in-situ crosslinking can therefore enable an increase in the selection of
inks for inkjet printing; among them are hyaluronan methacrylate, GelMA, polyethylene
glycol diacrylate (PEGDA) and norbornene-functionalized HA (NorHA) [161].
2.3. Stereolithography (SLA)
Stereolithography (SLA) is a 3D printing method based on polymerization in a mold.
In this process, light-sensitive fluid resin is dispensed into a mold (or container) and se-
lectively polymerized (i.e., hardened and solidified) by exposure to UV light (Figure 3).
Most resin 3D printers work that way, but a few “top-down” systems exist as well. In a
top-down setup, the light source sits above the resin tank in a top-down setup, therefore
curing the surface instead of the bottom. The build plate moves down to leave room for
new layers atop the previous ones until the object is complete and appears upright. The
UV light hardens the resin layer by layer so that the pieces can be built up in layers to
create the final object. In SLA, layer height (layer thickness) is usually about 50 µm but
may also be as little as 10 µm. Typically, the thinner the layers, the greater the quality and
the increased time is required for printing. Although SLA 3D printing is extremely versa-
tile and precise and creates smooth surfaces, a disadvantage of the method is that the parts
can degrade when subjected to sunlight over time due to the photosensitive characteristics
of the 3D printing resins.
Once the entire part is finished, it is removed from the tank and placed in a solvent-
based chemical bath to remove excess material and create a smooth surface. Finally, the
workpiece is hardened in an ultraviolet furnace to ensure that it is properly strengthened.
Figure 3. Schematic drawing of a SLA device.
The functional principle of SLA bioprinting works as follows. SLA represents an ini-
tial and original technique of 3D printing that forms the basis of modern 3D bioprinting.
In 1984, the first setup has been patented, and four years later the first commercial device
Printed construct
Biomaterial ink
or bioink tank
Light source
Light beam
Mirror
Cells 2024, 13, 1638 10 of 87
has been generated [162]. It is characterized by an extremely versatile choice of materials
and features the highest resolution and accuracy, as well as fine details. It is ideally suited
for functional prototyping. The SLA process originally contributed to establishing 3D
printing as a feasible option for manufacturers and inspired others to explore new print-
ing techniques and new uses for the technology. During the construction process, support
structures must be added so that the overhangs have some support. 3D printing with resin
also requires additional finishing steps, such as washing out of residual resin, breaking
away the support elements and post-curing, which is required to subject the printed ele-
ment to extra UV light for additional hardening.
What types of SLA 3D printing are there? The acronym “SLA” usually refers to ste-
reolithography, and in its original form, the light source used in resin 3D printing came
from lasers reflected by mirrors. Laser printers work very accurately, but they are very
costly to purchase and keep in good condition. The stereolithography of the present day
also includes other technologies such as digital light processing (DLP) and masked SLA
(MSLA). In DLP, a projector is used as the light source in place of a laser. The beam pro-
jector contains a digital micromirror device (DMD), where a micromirror constitutes a
pixel. The DMD is coupled with a visible LED, laser, or lamp light source for illumination
and can also employ UV or IR wavelengths for light-steering purposes. The light source
and the micro-mirrors work in synchronization to provide the desired projected output.
DMD controls the angle of each mirror and determines the brightness of the light trans-
mitted through the mirror. The DMD can therefore regulate the light intensity within a
small part of the projected beam. In DMD, the intensity of the pixels is color-coded. The
light beam produced from the DMD travels through a system of lenses and is focalized
on the pre-polymer hydrogel suspension. In zones exposed to strong light, the photoiniti-
ator captures enough photons to initiate polymerization, whereas in zones exposed to lit-
tle or no light at all, no polymerization takes place. Thereby, spatial crosslinking and SLA
bioprinting are implemented [163]. A new visible light-crosslinkable bioink that is de-
signed on the basis of cell-adhesive gelatin has been introduced [163]. The bioink com-
prises a photoinitiator derived from eosin Y (EY) and a GelMA prepolymer suspension,
which is suitable for the printing of cells and organoids. While laser SLA printers apply
layers of print dot by dot, DLP printers harden each layer at once in a single pulse of light.
This process is quite fast compared to conventional laser-based resin 3D printers. Similar
to DLP, MSLA hardens all complete layers at once. MSLA printers, however, use a series
of LEDs as a light source in place of a projector. The LED lights illuminate through an
LCD screen that selectively blocks the light by brightening or turning off certain pixels.
The resolution of an MSLA printer is therefore determined by the resolution of its LCD
display. SLA-based bioprinting offers benefits in terms of resolution and short printing
time, which is why it is currently attracting a lot of attention in the printing of cells and
organoids. Traditional SLA bioprinting, however, involves the use of UV light as a pho-
topolymerization principle, which can cause mutagenesis and carcinogenesis of cells
[163].
2.4. Laser-Induced Forward Transfer (LIFT)
In 1988, a type of laser-assisted printing technology was presented, the laser-induced
forward transfer (LIFT) (Figure 4) [164]. In 2004, LIFT was initially used for bioprinting,
whereby cell patterns were printed with excellent cell viability [165]. Thus, the emphasis
was placed on the LIFT technique, as it is highly useful to design 3D organoids and 3D
tumoroids. Unlike inkjet and extrusion printing processes, LIFT technology is character-
ized by high printing precision and high resolution, which is down to the micrometer
range, high output and a high rate of cell survivability [166].
Cells 2024, 13, 1638 11 of 87
Figure 4. Schematic drawing of LIFT device.
As no nozzles are required during the printing procedure, there is no problem with
ink blockage while printing. In addition, this technique can be integrated with different
bioprinting methods to broaden printing possibilities and offers the opportunity for in
situ printing [167]. As a consequence, LIFT has already been extensively exploited for the
bioprinting of pharmaceuticals [168], DNA [169], proteins [170], human osteosarcoma
cells [171], human endothelial cells [172,173] and mesenchymal stem cells [174] with phar-
maceutical drug administration and screening capabilities, nucleic acid microarrays, pro-
tein microarrays and printing of living cells, tissues and entire organs. The high cost of
LIFT technology is a major issue restricting its exploration and commercial deployment,
but this may change quickly due to high demand. The different variables associated with
LIFT bioprinting, comprising laser energy, laser spot size, physical characteristics of the
bioink and absorbing layer height, are discussed for efficient and successful bioprinting.
The principle of the LIFT process is to focus a light beam passing through a transparent
substrate onto a metal or polymer thin sheet, where part of the light is being absorbed and
transferred into internal energy [107,175]. This process raises the temperature, stretches
and distorts, and can even lead to fluidization or vaporization, resulting in the transfer of
material [176,177]. The LIFT system primarily comprises a laser unit, a donor with multi-
ple sheets and a recipient medium. The laser unit is usually a pulsed single-mode laser of
a specific wavelength.
The LIFT bioprinting donor layer typically comprises three elements: the substrate,
the absorbing coating and the biomaterial sheet [178]. Transparent glass is typically uti-
lized as a substrate for laser wavelengths in the near infrared or visible spectrum, while
quartz and fused silica are being employed for ultraviolet wavelengths. In addition, flex-
ible organic carriers are being investigated as possible substrates. The donor substrate,
such as transparent quartz glass with virtually no laser light absorption, a metal or metal
oxide-coated laser absorbing film and a coating of a biological solvent comprising biolog-
ical matter like DNA, proteins, or cells. The biomaterial film acts as an ink, which means
that it is printed, and the cell ink is made up of cells, cell culture fluid and matrix material.
The ink necessitates biochemical characteristics akin to those of the native ECM and usu-
ally comprises cell culture fluid [179], fibrinogen [180] or glycerol [181]. The matrix mate-
rial needs to be very close to the architecture and formulation of the ECM and should have
outstanding biocompatibility, moldability, minimal cell damage and be easily decompos-
able. An absorber layer is usually placed between the transparent substrate and the donor
film to avoid direct laser interference with the material to be deposited. Titanium (Ti),
titanium dioxide (TiO2) or gold (Au) is usually utilized as the metallic absorber coating
Printed construct
Light source
(Laser)
Beam
Droplet
Printed droplet
Focusing lens
Bubble
Laser focal
point
Absorbing
layer
Bioink layer
Target
slide
Target
slide
Donor
slide
Mirror
Cells 2024, 13, 1638 12 of 87
[165,182,183]. Different kinds of UV-absorbing films, such as polymer coatings [177], have
also demonstrated similar printing outcomes. The recipient substrate, generally a co-
verslip, is placed parallel to the donor and both are mounted on a moving 3D stage to
collect printed droplets [176]. The glass substrate is covered with a hydrogel coating or
other biocompatible substrate that is key to keeping the biomaterial vital following print-
ing [184]. Primarily, the layer works like a buffer that efficiently minimizes the shear im-
pact damage of the printed biological material on the substrate. Secondly, the hydrogel
moisturizes the biomaterial and avoids the volatilization of tiny droplets on the receiving
surface. Thirdly, the collagen and laminin present in the hydrogel also ease the attachment
of the printed organisms to the surface and assist their ongoing differentiation. For exam-
ple, the thickness of the layer on the capture substrate has been found to influence the cell
activity when printing multipotent embryonic cancer cells [165]. Without any buffer sub-
stance, the survival rate of the printed cells lay at 5%. When the coating height has been
raised from 20 to 40 μm, the cell survival rate increased from 50% to more than 95%. Nev-
ertheless, the ideal thickness of the layer varies according to the experimental setup, in-
cluding the viscosity of the bioink, energy of the laser and size of the dot. LIFT is based
on the concept of light–matter interference, where a portion of the light is absorbed by a
metal or polymer layer of the donor and converted into internal energy, resulting in a
temperature rise and blister generation when a laser beam is centered on the layer through
a transparent medium [176,183]. The bubbles then expand and deform, leading to their
collapse and the formation of a jet or droplet of bioink, which, in turn, enables mass trans-
fer and ultimately the printing of cells or other biological substances within the bioink
[177,178]. The receiving surface is a slide coated with buffer. As a heating device in LIFT,
pulsed laser systems with pulse widths of a few nanoseconds constitute the majority of
laser systems, even though ultrafast laser devices emitting picosecond and femtosecond
pulses may also be utilized [177,185]. Optical components like beam splitters and lenses
are utilized to manipulate, control and concentrate the laser beam onto the intersection of
the donor substrate and the layer of donor material. The laser wavelength must be ad-
justed to the transparency of the donor substrate and the absorbing ability of the donor
layer, even if it does not necessarily affect the process. In addition, the characteristics of
the laser system like laser energy density, pulse length, frequency and pulse energy have
a considerable influence on the process and the outcome. The selection of the wavelength
varies according to the interfacing material (interlayer or material to be deposited), with
ultraviolet radiation commonly chosen. With the LIFT bioprinting technique, the laser
pulse energy and the beam size are critical variables, whereas a number of other studies
have also highlighted the laser fluence, which is proportional to the pulse energy divided
by the spot size, as a pivotal factor [186].
To understand the theoretical nature of the material transfer mechanism at LIFT, nu-
merical analysis and simulation techniques have been used to analyze the heat generation,
thermal propagation and material transfer characteristics of various materials at a range
of laser energies and pulse durations [187–189]. Several theoretical models have been ar-
ticulated, comprising explosive ejection, phase alternation ejection and shock wave ejec-
tion. Firstly, the explosive ejection theory assumes that mass transfer is driven as a result
of the pressure produced during the laser ablation and vaporization, thereby causing an
explosive event [190]. When the melting boundary surface has not yet attained the air
boundary layer, the material has eroded and vaporized, and the gas pressure ejects the
material in an explosive manner in a very confined area. Secondly, the theory of phase
transition ejection can clearly provide an understanding of the ejection of metal micro-
droplets [191]. Based on this hypothesis, metallic materials stretch under laser irradiation;
however, they remain in a solid phase. Simultaneously, the focused laser forces the melted
boundary to progressively advance along the metal film in the direction of the air bound-
ary until the film–air border is also liquefied and the fully fluidized film ceases to be
trapped at the interface, creating a metal droplet that is discharged and transmitted.
Cells 2024, 13, 1638 13 of 87
Thirdly, the theory of shock waves says that the coating partly melts and volatilizes when
heated and simultaneously spreads in the direction of the substrate [192].
Multiple parameters influence a number of important factors including blistering, jet
evolution, deposition volume, resolution and cell survival throughout the LIFT bioprint-
ing procedure [176,183], among them firstly, the energy of the laser; secondly, the diame-
ter of the laser spot; thirdly, the physical characteristics of the bioink and fourthly, the
absorptive layer height. The generation of the beam during LIFT bioprinting involves
three distinct modes that vary with the level of laser pulse energy, such as the sub-thresh-
old mode, the jetting mode and the plume mode [193]. Below the threshold range, the jet
is unable to fully develop because of the lack of laser energy or excessive fluid viscosity,
which leads to a failure of material transfer. On the contrary, in plume mode, excessive
laser energy or insufficient fluid density can induce an unstable jet, which results in the
creation of unintended plumes and discontinuous droplets with different volumes. A sta-
ble beam that facilitates efficacious and regulated bioink release appears just when the
laser energy lies in between the beam and plume threshold values. The size of the laser
dot constitutes an additional key factor that influences LIFT bioprinting and defines the
printing resolution [166,185]. A narrower laser spot generally results in a better resolution;
however, it also carries considerably diminished ink when printing and, hence, leads to
less productive printing.
Viscosity is not just an key performance marker for the bioink, but it also has an im-
portant part to perform in bioprinting [189,194]. In case the viscosity of the bioink becomes
insufficient, splashing can arise while printing. Conversely, when the viscosity is exces-
sive, the laser needs more energy to initiate the inkjet printing process. A suitable viscosity
of the ink is essential to ensure a stabilized jet. Like the size of the laser dot, the viscosity
of the bioink affects the print resolution considerably, which is affected to a greater extent
than the laser energy, particularly in the case of decreased laser energy. The existence of
cells in the bioink is critical as it can substantially impact the LIFT bioprinting performance
[183]. In comparison to printing using non-cellular bioink, cell entrapment usually needs
increased laser energy, leading to lower beam velocities and narrower printing dots [195].
Moreover, the uneven dispersion of cells caused because of the aggregation within the
bioink can generate two kinds of non-ideal beam characteristics while printing, such as
non-straight beams harboring non-straight trajectories and straight beams harboring non-
straight trajectories. As explained above, the depth of the absorber layer and the gap be-
tween the donor and receiver layers can affect the success of the printing procedure. Fi-
nally, when applying LIFT for laser bioprinting, it is essential to set suitable parameters.
Consequently, this may contribute to keeping up the printing pace and prevent biological
alterations, like deterioration of phenotypic or nucleic acid cell integrity during LIFT bi-
oprinting. LIFT technology has already been used to print a range of other biomaterials,
including lipid vesicles for drug delivery and biosensing applications. Lipid vesicles,
however, are molecular partitions made up of lipid bilayers and can be tens of microme-
ters thick, complicating printing with conventional direct printing methods. As the LIFT
technology enables the printing of objects with large dimensions, lipid vesicles have been
successfully printed with LIFT without compromising the vesicle membrane [196]. The
printing of vascular structures, in contrast, is still in its infancy [173] and requires signifi-
cant improvements regarding the vessel stability, the vessel modification and its function-
ality [197].
3. Usage of Several Bioinks and Bioink-Database for Cells, Hydrogels and Organoids
The complexity of 3D bioprinting is enormous due to the large amount of different
bioinks available. This section describes basically the major issues of bioinks in general
from a biophysical cell perspective. There are detailed review articles available for more
background and refined information [198,199]. It is therefore natural that the overview
and usability of bioinks has been summarized in a first bioink database for 3D extrusion
printing [200]. The database enables the easy identification of combinations of extrusion
Cells 2024, 13, 1638 14 of 87
pressure, temperature and speed that have been optimized for the printing of specific bi-
omaterials and, even more importantly, to highlight the areas in which printing cannot be
accomplished. The database allows scientists and prospective bioprinting users to quickly
find the right bionics for the respective application and helps with the execution of the
printing by utilizing decisive parameters that must be considered in each case. This data-
base is constantly being expanded through the voluntary input of new bioinks and their
printing parameters. The collected results also permitted a correlation analysis among all
printing variables, such as needle size and type, that showed suitability for cell-based 3D
printing. Although bioprinting is still in its infancy, the important issues of standardiza-
tion and evaluation of factors, such as the shape accuracy of the printed structures, repeat-
ability, material characteristics and the used hardware and software, have been purposely
addressed from a regulatory and clinical viewpoint [201–203]. However, that does not
account for the vast variation in printing regimes, which are specific to every printing
mode and 3D printer utilized and reported by researchers worldwide [204]. As a result,
there is a wealth of data that is both useful and sometimes confusing and contradictory in
the 3D bioprinting field. Therefore, the establishment of the world’s first bioink repository
for 3D bioprinting is a logical consequence to make it easier to navigate and keep track of
things. The database is open source and enables researchers to simply access it and add
the results of their work to the database repository. This database concentrates exclusively
on microextrusion printing and captures critical printing variables like the composition of
the bioink, pressure, temperature, velocity, needle type and the cell type employed. The
database is freely accessible at https://cect.umd.edu/3d-printing-database (accessed on 20
August 2024). At present, there are more than 200 various bioink compounds listed that
have been utilized for 3D printing. These materials comprise thermoplastics like PCL,
PLA, PLGA, natural and synthetic polymers like alginate, collagen, decellularized ECM
(dECMs), and PEGs, ceramics comprising hydroxyapatite and β-tri-calcium phosphate,
various nanomaterials, nanocomposites, biomolecules and proteins, which have been uti-
lized as additional ingredients in the extruded bioinks. The following are the physiochem-
ical properties of the selected polymeric hydrogels that serve as bioinks are briefly pre-
sented. As the hydrogels serve as a scaffold for the printed organoids, apart from bio-
chemical and structural characteristics, the mechanical characteristics seem to be im-
portant for organoid survival, growth and further development. In the following, the dif-
ferent types of hydrogels employed for bioprinting are presented. Beyond the mechanical
properties of hydrogel scaffolds, there are degradable and non-degradable scaffolds,
which may be relevant in processes, where temporal stability is required, but need to be
altered over time to mimic physiological conditions.
3.1. Physiochemical Characteristics of Polymeric Hydrogels Employed for Cells and Organoids
When using inks that are compatible with living organisms (referred to as bioinks),
non-toxicity, degradability, cell adhesion and porosity must be guaranteed [62,198,205–
208]. Inks in which living cells are encapsulated are in a state of conflict, as the properties
that constitute a stable printing, such as density or viscosity, are frequently in direct op-
position to the maintenance of viability, as cells require a porous and compliant surround-
ing in order to grow and migrate [209]. The rheological demands on the bioink change
depending on the bioprinting method, such as inkjet or droplet-based, laser-based or ex-
trusion-based printing [201,207,209,210]. In inkjet bioprinting, a continuous stream of
small droplets is used to generate the 3D structure. This process is, nevertheless, generally
time-consuming and inefficient for the production of tissue on a large (clinical) scale [206].
Laser-based bioprinting utilizes a precise laser beam to harden the engineered structure
in a pool of bioink, but heat can harm the cells [210] and the process is quite slow [201].
Extrusion-based bioprinting comprises the shape of a low-viscosity filament or thread
during printing, which hardens on the print surface, retains its shape and encourages the
layering process [206]. A major obstacle with this technique is that the cells are subjected
to a perceptible shear stress from the applied extrusion pressure as they travel through
Cells 2024, 13, 1638 15 of 87
the syringe and nozzle, potentially leading to cell injury [201,208,211]. To mitigate this
stress, the bioink must exhibit lower viscous properties [212], but this can lead to distor-
tion, collapse and occlusion of pores, which, in turn, reduces accuracy and resolution
[209]. Hydrogel-based bioink compositions are a versatile choice for a wide range of tech-
niques to accomplish bioprinting. Polymeric hydrogels consist of 3D interconnected scaf-
folds of hydrophilic polymer chains that can absorb significant quantities of water and
expand up to 99% of their dry weight in water (w/w) without disintegration [213,214].
Although there are different kinds of hydrogels and different gelling techniques, the em-
phasis, in this review, is placed mainly on polymeric hydrogels. Polymeric hydrogels offer
excellent biocompatibility and tissue-like mechanical characteristics, rendering them ideal
for 3D bioprinting and a range of tissue engineering uses [49]. Hydrogels mimic the ECM,
the natural environment of cells, in an effective manner and offer a hydrated and textur-
ally supportive surrounding that can be populated with cells homogeneously and in an
efficient way [214]. Cells can be dispersed in these polymeric hydrogels to produce a bio-
ink and in a controlled manner in bioprinting applications. Cells can be embedded into a
hydrogel emulsion, and this bioink has been extruded to form cell-laden vascular patterns
[49]. Hydrogels frequently consist of shear-thinning materials that can be forced into ex-
trusion under high shear stress and subsequently retain their mechanical characteristics
[214,215]. Therefore, they are ideal for bioprinting purposes. Substances like gelatin, PEG
and Pluronic® behave like fluids during the printing process and return to a gel-like con-
sistency after the extrusion, ensuring that the printing process has the required stability
to create the intended texture. The following subsections outline the characteristics of nat-
ural and synthetic polymer hydrogels typically encountered in bioprinting processes.
Knowledge of the physicochemical characteristics of hydrogels is essential for evaluating
the stability, performance and toxicity of hydrogel uses in bioprinting. The most relevant
physicochemical characteristics are the pH value, the printing temperature, and the de-
gree and type of crosslinking.
3.1.1. pH
The majority of hydrogels can be stored and processed via bioprinting under physi-
ological pH conditions, such as around 7.4 [216,217]. The pH value of hydrogels signifi-
cantly affects the swelling properties of hydrogels [217]. The swelling capacity determines
the form and volumetric variations of a hydrogel; thus, a higher swelling ability is favored
because of the enhanced robustness of the hydrogel [218]. The highest swelling potential
of most hydrogels is at a physiological pH value of approximately 7.4 compared to an
acidic or basic pH value [217]. A shift in pH generally leads to a modification of the poly-
mer chain charge, resulting in either swelling or non-swelling of the hydrogel and a gen-
eral modification of stability [219]. In particular, pH-sensitive hydrogels are prone to pH
variations, mainly due to their ionic character [219]. At low pH, cationic hydrogels have a
natural propensity to swell because of the protonation of amino/imine chains, while ani-
onic hydrogels tend to swell at higher pH values because of the ionization of acid chains
[219]. Measuring the swelling ratio can also yield insights into the type, amount and tight-
ness of crosslinking in the polymer matrix and can be utilized to indirectly assess the me-
chanical characteristics of the gel, like the modulus of elasticity (E) (stiffness) [220,221]. A
higher pH value of collagen during the gelation process leads to higher stiffness [222],
which can reduce the printability and the vitality of cells during bioprinting. When the
pH value was raised in the region of 5 to 8, for example, the relaxation modulus of collagen
gels increased linearly (in other words, the gels stiffened) and then stagnated. At the same
time, the viscosity of the hydrogel changes, such as a low viscosity at higher pH of 8.5
[223], which also affects the shear thinning behavior that is critical for bioprinting. In ad-
dition, the pH value affects the gelling time, which impacts the overall cell survival and
proliferation.
Cells 2024, 13, 1638 16 of 87
3.1.2. Temperature
The temperature is inversely related to the viscosity of the hydrogel [224]. The higher
the ambient temperature, the lower the viscosity, which is associated with reduced shear
stress and minor deterioration of the cells [224]. For bioprinting, a low viscosity of the
bioink is required to achieve an optimal printing result in terms of cell viability, but there
is often the problem of suboptimal print accuracy and image resolution [225]. However,
in bioprinting, the printing temperature required varies according to the type of polymer
utilized. In reactive ionotropic polymer printing, the polymer liquid can be kept and
printed at cell culture temperature, such as 37 °C, for the production of hydrogels [226].
Since gelation in the reactive printing of ionotropic polymers is initiated in a tank with
suitable counterions, gelation is very fast and it is possible to print polymer solutions to-
gether with cell culture media [226]. There is a lag phase during collagen gelation in which
the primary aggregates of collagen molecules are established (nucleation event). Next, mi-
crofibrillar aggregation begins with the lateral aggregation of subunits triggered by alter-
ations in ionic strength and pH and increases the temperature up to 37 °C until reaching
equilibrium. In opposition, the fundamental mechanism of gelatin is associated with the
reverse coil-to-helix transition induced when solutions are cooled below 30 °C, with the
resulting helices resembling the collagen triple helix but not achieving equilibrium. The
gelling processes are thermoreversible for both collagen and gelatin, albeit in opposite
directions: collagen gels dissolve when the temperature is decreased, whereas gelatin gels
dissolve when the temperature is elevated [227].
For hydrogels made up of polymers that react through physical interactions, the op-
timal temperature depends on the type of polymer to be gelled. Hydrogel materials, like
gelatine methacryloyl, often experience a physical sol-gel or gel-sol transition from room
to body temperature and can also be chemically crosslinked at these temperatures to
achieve dimensional stability [228]. Normally, heated polymer mixtures are printed in a
chilled surrounding, in which they attain their gel transition temperature and solidify
[226]. Agarose, for instance, is soluble in water at above 65 °C and melts into gel at 85 °C
[229]. For this reason, agarose is usually placed in the printer tank at temperatures ranging
from 60 to 80 °C [226]. The agarose is subsequently printed in a chilled liquid bath, typi-
cally between 17 and 40 °C, which is lower than the gel transition temperature [199]. For
polymers gelling by physical contact, the final gel temperature frequently impedes cell
embedment and shortly following printing, as the temperature is outside the normal body
temperature regime and is potentially detrimental to living cells. Moreover, in depend-
ence to temperature the some materials exhibit a shape memory effect [230] indicating
that the temperature is highly relevant.
3.1.3. Crosslinking
The majority of bioinks, such as hydrogels, are stabilized by crosslinking mechanisms
to maintain the shape and mechanical strength of the 3D-printed structure. Post-print
crosslinking is a procedure in which the interior architecture of the printed hydrogel is
altered to preserve the overall structural integrity and obtain the mechanical characteris-
tics of the bioprinted structure [224]. The two commonly occurring crosslinking mecha-
nisms are of physical and chemical nature [109]. Physical crosslinking is achieved through
physical processes, and comprises intermolecular interferences between polymer chains
like hydrophobic interference, electrostatic interference, hydrogen bonds, stereocom-
plexes and guest–host interference [109]. The physical crosslinking phenomenon can be
reversed, and few or no chemical responses are required to create this connection. In
chemical crosslinking, reactants are applied to effect the covalent bonding of chemically
responsive functional chains [109]. Chemical crosslinking is generally mechanically more
powerful than physical crosslinking, as it produces covalent bonds between the polymer
chains, but, whichever agent is applied, it carries the potential risk of causing cytotoxicity.
The chemical crosslinking of hydrogels has an irreversible effect; however, its benefits are
Cells 2024, 13, 1638 17 of 87
durability, adjustable structures and excellent mechanical characteristics [109]. In enzy-
matic cross-linking, covalent bonds are also established among the polymer chains, but
the extent of crosslinking is somewhat less strictly regulated [198]. Managing the level of
crosslinking is critical in bio-implementations as it allows the stiffness of the structure to
be altered and adapted to the respective tissue [231]. A less common cross-linking phe-
nomenon is thermal cross-linking, a mechanism that requires temperature fluctuations in
the environment. The majority of natural polymer hydrogels crosslink at 37 °C [109]. Nev-
ertheless, only a few hydrogels, such as alginate and gelatin, crosslink at room tempera-
ture [109]. When a gelatin suspension begins to cool, the protein polymers coil into
twisted, helical configurations, causing the mixture to harden. Other natural and synthetic
hydrogel substances display a comparable temperature sensitivity, including gellan gum
(that can also be ionically crosslinked), agarose, polymers of N-isopropylacrylamide
(NiPAAM) and Pluronic F127 (synonymously referred to as Poloaxomer 407). The sol-gel
transition in the product Pluronic F127 takes place upon heating, which means that the
polymer solution forms a liquid state at low temperatures and becomes a thermally cross-
linked hydrogel when heated [232]. The viscosity and stability of Pluronic F127 can be
improved by adding chitosan [232].
3.2. Biological Characteristics of Hydrogels Serving as Bioinks for Cells and Organoids
Hydrogels have the utility for bioprinting as they exhibit numerous characteristics
akin to those natural extracellular matrix of tissues, facilitating the embodiment of cells in
a highly hydrated form and mechanically maintain the 3D environment [134]. The bionic
scaffold should have adequate mechanical strength, biocompatibility, cell proliferation,
survival and other biological properties. The drawbacks of hydrogels made from natural
polymers are poor mechanical characteristics and low printing performance and dimen-
sional stability. In recent years, a number of synthetic, modified and nanocomposite hy-
drogels have been designed that can modify their characteristics by physical interactions,
chemical covalent bond crosslinking and bioconjugation reactions to meet the specifica-
tions [233]. The hydrophilicity of hydrogels constitutes the primordial driver of biocom-
patibility, rendering them a beneficial channel for the production of tissue structures [234].
Hydrogels offer a proper microenvironment for cell proliferation and are highly adapta-
ble. They enable a range of biochemical and biophysical features to regulate cell behaviors,
among them cell adhesion, migration, proliferation and differentiation [235]. A variety of
cell types are capable of survival when embedded in hydrogels, as these scaffolds can
provide cell–matrix interactions to determine cell fate [236]. Among these cell types are
fibroblasts, chondrocytes, macrophages, hepatocytes, endothelial cells, smooth muscle
cells, adipocytes and stem cells [134,236]. There is a complicated communication between
hydrogels and cells, including stem cells; multiple parameters like porosity, different
types of polymers, stiffness, tunable heterogeneous structure via magnetic beads, compat-
ibility and decomposition cause the survival or death of stem cells [237–240]. Hydrogels
imitate the 3D ECM and create a favorable environment for cells. Cells, including cancer
cells and stem cells, can perceive their environment via mechanosensing through various
elements, such as cell surface receptors, such as 𝛼5𝛽1, 𝛼𝑣𝛽3, 𝛼1𝛽1 or 𝛼2𝛽1 integrins,
caveoli, ion channels, such as Piezo 1 and Piezo 2 and extracellular vesicles, including
exosomes, in order to further develop, expand, multiply (proliferate) or simply survive
[241–246]. Hydrogels can be generally made of pure natural or pure synthetic polymers
or mixtures of both [199]. Both natural and synthetic substances with various characteris-
tics and performances are utilized in the manufacture of hydrogels [224]. Synthetic hydro-
gels have increasingly been employed more frequently than natural polymers in recent
times due to their increased water absorbency, extended durability and the variety of
available chemical ingredients [199]. The natural and synthetic hydrogels are presented
below, and some examples of each type are explained in more depth. Moreover, hydrogel
compositions and their adaptations are characterized. Modifications enable changes in the
chemical performance and mechanical characteristics of hydrogels [247]. For synthetic
Cells 2024, 13, 1638 18 of 87
hydrogels, changes are critical for enhancing biocompatibility and cellular attachment
characteristics, whereas for natural hydrogels, changes enhance shaping capabilities.
Chemical changes to hydrogels could contribute to the creation of robust hydrogel struc-
tures enhancing characteristics like dynamic coupling, shear thinning and self-healing
and promoting covalent as well as ionic crosslinking [199].
3.2.1. Natural Polymer Hydrogels Serving as Bioinks for Cells and Organoids
Natural polymers originate from natural material sources. Typically employed natu-
ral polymers comprise cellulose, collagen, guar gum, gelatin, chitosan, alginate and fibrin
[199,248]. Hydrogels made from natural polymers exhibit superior biological characteris-
tics compared to their synthetic equivalents, as they comprise improved biocompatibility,
biodegradability and procellularity [224]. The rationale for this is that natural polymers
can coat the surface of eukaryotic cells and then bind with proteins to form a natural ECM
[199]. For instance, glycosaminoglycans (GAGs), unbranched high molecular weight pol-
ysaccharides that are either covalently bound to protein cores and constitute proteogly-
cans or occur freely in the ECM, can coat the surface of cells and couple with several pro-
teins to produce a natural ECM, leading to outstanding biocompatibility and cell affinity
[249]. The integration of GAGs in biomaterials offers novel pathways for the display of
signaling molecules and facilitates the monitoring of development, homeostasis, inflam-
mation and the development and propagation of tumors. GAGs provide the structural
foundation for several important functional characteristics of the ECM. Besides the hydro-
gel characteristics of tissues, such as compressive resistance [250], GAGs convey the local
display of multiple soluble signaling molecules [251], involving the generation of mor-
phogen gradients [252]. For example, it has been demonstrated that HS-GAGs regulate
the generation of morphogen gradients in vivo [253], and thereby control the adaptive
development of tissues and organs within multicellular organisms [254]. Thus, GAGs
have been demonstrated to be associated with key events, including in development, and
tumor evolution and malignant progression [255,256]. Ultimately, GAG-containing hy-
drogels can be rendered vulnerable to breakdown by cell-secreted proteases by integrat-
ing matrix metalloprotease (MMP)-cleavable peptide crosslinkers [250]. In addition, the
majority of natural polymers possess bioactive constituents involved in amplifying extra-
cellular signal transduction to enhance cell proliferation, cell differentiation and cell func-
tionality [199]. These components comprise protein ligands and motifs that attach to cells;
thereby, forces can be transduced from the ambient environment toward the cell’s interior
[257]. Many of the following natural biomaterials can be employed for 3D organoid cul-
ture. Thus, they are briefly introduced in the following.
• Agarose
Agarose as a natural polysaccharide is derived from marine algae. It is not as com-
monly employed for bioprinting purposes as some other natural hydrogels because it is
challenging to print and, as it is extracted from a plant, it lacks biomimicry for mammalian
cell types [258]. Nevertheless, its beneficial gelling characteristics render agarose an at-
tractive hydrogel constituent and supporting framework. In nozzle-based bioprinting,
agarose made its first appearance in 2005, when Chinese hamster ovary (CHO) cells were
printed in agarose scaffolds [152]. In recent times, an agarose in a compound hydrogel,
consisting of gelatin and alginate, in which adipose-derived stem cells (ASC) are sus-
pended, was introduced [259]. Highly precise and robust bioprint textures were printed.
It has also been observed that the addition of agarose enhanced the pore size and quantity
in the hydrogel, favoring cell proliferation [259]. Other studies show that agarose can be
used effectively in a very indirect way. In 2018, Mirdamadi et al. reported a technique of
embedded bioprinting that built on the seminal research of Hinton et al. in 2015 [260], in
which a cell ink was expressed in an agarose suspension [261]. The agarose suspension
offered temperature-resistant textural enhancement to the soft bioprinted structures
throughout and beyond printing, so that the printed construct could stay in the
Cells 2024, 13, 1638 19 of 87
suspension even when transferred to the growth incubator. Moreover, the agarose gel was
penetrable for constituents of the cell medium, which resulted in media replacement
through diffusion in the vicinity of the printed structure with no disturbance of the texture
[261]. In 2016, the application of agarose in conjunction with collagen in a nozzle-based
bioprinter was published [262]. When agarose was added to collagen, a tissue-like matrix
was created. The advantage was that the addition of agarose to the cell ink did not alter
the structural topography of the collagen mesh and the collagen solution had no effect on
the agarose gelling. The incorporation of agarose into the cell ink resulted in a more vis-
cous ink, a reduced droplet size and increased printing precision [262]. In 2022, efforts in
extrusion-based bioprinting characterized the agarose-gelatin hydrogel mixtures charac-
terizing the mechanical and rheological characteristics for bioprinting [263]. Moreover, the
human SH-SYn5Y neuroblastoma cells from the neural crest [264] were printed using the
above-mentioned agarose-gelatin mixture as cell ink and differentiated into neuron-like
cells [263].
• Collagen
Collagen type I is among the top prevalent fibrous proteins in the ECM and is the
principal structural component of the ECM that offers tensile strength, controls cell adhe-
sion and aids in cell proliferation [265,266]. These properties render collagen to be an ideal
hydrogel for the application in cell inks for bioprinting, as multiple tissue cells can gener-
ally adhere to it [267]. The principal types of type 1 collagen are pig skin, rat tail tendon
and cow skin [268]. However, all these types of collagen can exhibit diverse mechanical
and structural cues upon scaffold formation [8]. The usage of collagen in cell inks, though,
is restricted because of its long gelling time, lasting up to 30 min at 37 °C [266]. In addition,
this long gel time can lead to an inhomogeneous arrangement of the cells and, conse-
quently, result in a loss of structural accuracy in the final printed object [266]. Moreover,
collagen is fluid at low temperatures and becomes fibrous at higher temperatures or at
neutral pH, which can be problematic when printing with nozzles, because the nozzle
mechanism is occasionally produced with heating [198]. In 2019, Lee et al. described the
application of freeform reversible embedding of suspended hydrogels (FRESH) for the
biological engineering of human heart parts at different levels of complexity [269]. In
FRESH, a collagen-based cell ink is extracted into a thermoreversible carrier pool consist-
ing of a suspension of gelatine microparticles, which acts as a carrier while the print is
being made and is then discarded. A left ventricle has been produced from human stem
cells with the help of the FRESH technique. In a two-material printing procedure, the col-
lagen ink and a cell ink with a high cell count density are deposited to generate the ven-
tricle. In the complete ventricle, it was possible to monitor synchronized contractions, the
directional propagation of the action potential and the wall-thickening characteristic of a
ventricle [269]. An aerosol jet bioprinting process has been developed for printing high-
density collagenous textures [270]. Aerosol jet bioprinting involves a printing process in
which an aerosol is generated from an ink and a vehicle gas that impregnates and coa-
lesces on a surface [270,271]. This technique may be an attractive way to print collagen
into high-density structures to use as a cell substrate [270], although several reports sug-
gest that high-density collagen structures can inhibit cell proliferation and impede the ca-
pacity for differentiation and diffusion of by-products [55,272]. Inversely, fibroblasts can
be cultured in high-density collagen gels (40 mg/mL) with a high viability rate after cul-
turing for a week [267], which highlights the opportunity to use aerosol jet bioprinting as
a novel tool to generate substrates for bioprinting.
• Fibrin
Fibrin refers to a fibrillar protein formed from fibrinogen that circulates in the blood
and is frequently derived from the plasma of mammals. A fibrin clot is the body’s first
response to a laceration as it builds a matrix of fibers to stop the bleeding. Fibrin, as uti-
lized in tissue engineering applications, is manufactured exactly as the body makes it by
activating fibrinogen monomers to form a polymeric fibrin matrix [258]. Fibrin is
Cells 2024, 13, 1638 20 of 87
degraded in the human body by fibrinolysis that is carried out predominantly by plasmin.
In vitro, cells generate enzymes that catabolize fibrin [273]. Therefore, fibrin hydrogels
suffer from a major deficiency in structural robustness for use in direct cell contact situa-
tions. Moreover, fibrin is a difficult choice for nozzle-based bioprinting due to its high
viscosity, and fibrinogen offers poor texture and form accuracy [273]. Therefore, fibrin
may be a difficult substrate in cell inks. A number of different approaches can be em-
ployed to overcome these restrictions and enable the effective incorporation of fibrin into
cell inks for bioprinting. A method is the utilization of fibrinogen, that has a viscosity close
to water. Once the fibrinogen has been deposited, the crosslinking reagent thrombin can
be given on to the fibrinogen or as a substrate to generate a definitive fibrin network
through the crosslinking of the fibrinogen through a calcium-dependent route [274,275].
This technique can be utilized with nozzle-based bioprinters that ink-print human micro-
vascular endothelial cells (HMVECs) with fibrin to produce a microvasculature [274]. The
use of an extrusion-based technique to print a fibrinogen-based cell ink in a thrombin-
enriched PEGDMA alginate pool has been reported to hyperlink the fibrin [273]. With this
technique of bioprinting, a soft microenvironment that mimics the soft pericellular matrix
of cartilage has been obtained. This enables improved nutrient delivery in a bioprinted
cartilage scaffold and thus the production of cartilage that is closely resembling that of
natural cartilage. A nozzle-based technique has been adopted to bioprint a three-layer
vessel wall for a vascular model [275]. Following printing of a surrogate gelatin core
loaded with human umbilical vein endothelial cells (HUVECs), a fibrin-based ink has
been bioprinted onto the gelatin core (lumen). The gelatin has subsequently been lysed,
and the retained ECs have been permitted to settle and adhere alongside the lumen of the
fibrin-based vascular graft [275]. In this manner, the adhesive-like characteristics of fibrin,
along with the strength and optimal surrounding for cells that it affords, could be har-
nessed. Although more advanced printing techniques are necessary than with traditional
hydrogels, the advantages of fibrin encompass its biological decomposability, its adhesive
characteristics, its adjustable mechanical and nanofibrous textural characteristics [273].
• Gelatin
Gelatin is another common component of hydrogels. Gelatine is often chemically
modified or mixed with another polymer before being processed into a hydrogel because
of gelatine’s inferior rheological characteristics [224]. In a number of trials, gelatin has
been altered with furfuryl chains to produce furfuryl gelatin (f-gelatin) [224]. f-gelatin can
be quickly crosslinked in the physical presence of visible light and retains its textural fi-
delity following crosslinking [276]. f-gelatin can also be amended with hyaluronic acid to
provide improved viscosity and shear thinning and to enhance the textural integrity and
rigidity of the reticulated structure [276]. Amending gelatin with free radical crosslinkable
methacrylic groups, which results in gelatin-methacryloyl (GM or GelMA), is a new tech-
nique to increase the stability of gelatin and enable its utilization in cell bioinks for bi-
oprinting and other tissue engineering applications [277]. The GelMa can be stabilized by
fluorenylmethoxycarbonyl diphenylalanine (Fmoc-FF) crosslinking peptide into the gel
bioink to overcome the post-printing processing of the bioink [278,279]. The effects of the
cooling and heating rates on sol-gel and gel-sol transitions in GelMA can be analyzed with
rheological techniques [228]. Crosslinking chemically modified gelatin at low tempera-
tures can lead to a higher modulus (stability) than the crosslinking carried out at high
temperatures [228]. The characteristics of the final hydrogel are thus highly sensitive to
the temperature of processing and can be adapted to the required use. By chemically mod-
ifying a gelatin-based hydrogel with glycidyl methacrylate, a protein-based elastic hydro-
gel (GELGYM) has been produced that can be specifically engineered for ocular tissue
engineering purposes; however, it can also be employed for various other tissue types
[280]. An engineered blood vessel could be developed to withstand a pressure of up to
350 mmHg, which therefore qualifies GELGYM as an attractive choice for a cell ink for
vascular bioprinting [280]. A mixture of methacryl-modified gelatin (GM), non-modified
Cells 2024, 13, 1638 21 of 87
gelatin and acetylated GM could be utilized to create vascularized osseous constructs
[277,281].
• Alginate
Alginate comes from brown algae and is a natural polysaccharide copolymer that is
among the natural polymers most frequently utilized for bioprinting [224,282]. As a bioink
for cells, it has multiple benefits as it is non-immunogenic, biodegradable, non-cytotoxic,
inexpensive and rapidly gellable [283]. The drawbacks are low cell adherence and the in-
sufficient promotion of cell proliferation [284]. In addition, alginate is difficult to print on,
and although it is a biodegradable material, alginate degradation can involve complicated
mechanisms. Alginate is hydrophilic and can therefore be blended readily with a series of
natural and synthetic polymeric cell inks, such as collagen [285], silk fibroin [286], and
decellularized and solubilized ECM (dECM) [287], to create a more favorable environ-
ment for cells compared to alginate on its own. The combination of these materials enables
a perfect balancing of biological and physical characteristics, with alginate frequently act-
ing as a textural stabilizer and as a thickening material. A chemical amendment to opti-
mize the characteristics of alginate is the oxidation of alginate. Oxidized alginate (ox-alg)
exhibits a quicker breakdown capacity and contains a higher number of reactive moieties,
thereby improving alginate’s suitability for sustaining cell performance [224]. Another
conventional alginate modified form is methacrylated alginate (MeAlg/AlgMA) [224].
Methacrylated alginate offers the capability of photocrosslinking, which opens more de-
sign possibilities for adapting the mechanical characteristics of the hydrogel, the pore size
scale and the decomposition velocity [288]. In addition to amending the hydrogels on their
own, new techniques are also being tried out to achieve better printing performance. A
combination of PEG and alginate leads to very long-lasting and extensible hydrogels [289].
The printed structures are made particularly durable by the inclusion of nanoclay. Micro-
structured alginate hydrogels have been produced by a microreactive inkjet printing
method in which a precursor and a crosslinking agent encounter each other in air while
printing [290]. This novel technique offers a unique option for jet-based bioprinting and
demonstrates favorable characteristics of the bioprinted alginate [291]. Typically, alginate
bioprinting is performed by one of two techniques: the alginate is pressed into a bath of
crosslinker, such as typically calcium, or the crosslinker is printed onto the precipitated
alginate [290]. This technique enables a freestanding vessel system with a small circum-
ference to be printed.
• Hyaluronic acid (HA)
Hyaluronic acid (HA) is a straight polysaccharide that occurs naturally in the ECM
of both cartilage and joint synovial fluid [292]. HA acts to preserve the synovial fluid by
enhancing its viscosity and increasing the flexibility of the cartilage. Thus, HA is extremely
biocompatible and promotes cell signal transmission, the healing of injuries and the or-
ganization of the matrix [293]. In addition, HA has been found to possess anti-inflamma-
tory properties, rendering it an attractive material for the implantation of bioprinted tex-
tures [294,295]. HA is negatively charged, which causes the attraction of cations and os-
mosis to absorb water, forming a gel [258]. Nevertheless, HA is easily soluble at room
temperature, which restricts its textural integrity and stability. HA has the potential to be
chemically altered with a range of functional chains to reduce breakdown and improve
durability [293]. A thiol-modified hyaluronic acid and thiol-modified collagen hydrogel
has been reported to be suitable for printing using a nozzle-based (jet-based) bioprinter
[221,296]. A drawback is that substantial dilution and chilling is required to jet this hydro-
gel substrate correctly while avoiding obstruction of the printer nozzles. Although di-
luted, this hydrogel readily underwent cross-linking at room temperature and offered a
sustaining medium for downstream cell ejection. Alginate-hyaluronic acid hydrogels can
be networked by multiple mechanisms, which include acyl-hydrazone, hydrazide inter-
actions and calcium ions [297]. It has been possible to prepare an alginate acyl hydra-
zide:HA monoaldehyde gel with a ratio of 50:50 (A5H5), with a gelation time of
Cells 2024, 13, 1638 22 of 87
approximately 60 s, a viscosity of approximately 400 Pa at a zero shear rate, high resistance
to different pH solutions and a prolonged breakdown time of over 50 days [297]. Moreo-
ver, intricate patterns like small, empty cylinders could be printed with no difficulty. In
2019, the bioprinting of skeletal grid structures with a cell ink comprising HA, hydroxy-
ethyl acrylate (HEA) and gelatin methacryloyl has been presented [298]. Moreover, stable
rheological characteristics and outstanding biocompatibility have been found [299].
• MatrigelTM
MatrigelTM stands for the trade name for the basement membrane matrix obtained
from the Engelbreth–Holm–Swarm (EHS) mouse tumor (sarcoma). The MatrigelTM is a
mixture of proteins and small molecules, mainly collagen IV, perlecan, laminin and
growth factors, and closely replicates the extracellular environment of many types of tis-
sues [258]. Matrigel is usually stored at 4 °C (liquid), and it undergoes polymerization at
the body temperature of 37 °C [258]. This property has made this hydrogel an outstanding
choice for bioprinting purposes. It is frequently employed in cell cultures as it potently
stimulates cell proliferation and cell differentiation. Cells grown on a MatrigelTM display
show a complicated cellular response that is usually hard to stimulate in a laboratory
[300]. Bioprinting has been performed with pure MatrigelTM suspensions containing hu-
man skeletal muscle progenitor cells [301], using a chilled print head to suppress gelling
of the hydrogel during printing and only allowing it to gel when deposited on the printing
surface at room temperature. After culturing the printed structures, skeletal muscle tissue
emerged that contained contractile, cross-striated myofibers that contracted in response
to electrical impulse activation. This type of bioprinted microphysiological system (MPS)
is beneficial for drug discovery, for instance, when testing drug candidates to treat muscle
atrophy [258]. A customized extrusion bioprinter has been utilized for the bioprinting of
mouse prostate cancer cells floating in MatrigelTM [302]. A volumetric dosing system has
been implemented to ensure that the irregular “splashing” extrusion that can arise when
printing plain MatrigelTM is minimalized. Although MatrigelTM has favorable characteris-
tics in terms of cell proliferation, it needs some adjustments concerning its printability
[303]. As an alternative for MatrigelTM, a biosafe dECM can be employed [303]. A major
weakness of MatrigelTM is the batch to batch variability [304].
3.2.2. Synthetic Polymer Hydrogels
Synthetic polymers are generally grouped into plastics, elastomers, and synthetic fi-
bers [305]. For tissue engineering, it is ideal to imitate the ECM to produce an optimal
tissue equivalent. Although synthetic hydrogels offer the benefit of photopolymerizability
and a high degree of adaptable mechanical characteristics, they cannot mimic the ECM
because of their bioinert nature [306]. Synthetic hydrogels are more water-absorbent com-
pared to natural hydrogels. The proportion of water in the hydrogel is dictated on the
basis of the characteristics of the polymer and the crosslinking density [307]. The simula-
tion of the ECM is essential because the ECM is not only a structural scaffold, but also
regulates cellular functions such as cell migration, cell proliferation and cell differentia-
tion [308]. Mimetic modification of the ECM in synthetic hydrogels has been shown to be
an effective means of eliciting the intended cellular reactions. Synthetic polymers create
artificial environments [224], whereby plastics, elastomers and synthetic fibers are the
most frequently used raw materials for the creation of synthetic hydrogels. Synthetic hy-
drogels can be easily manufactured and chemically modified for specific applications
[309]. Hydrogels made of natural polymers were initially increasingly debated due to their
favorable biological characteristics and later rejected in favor of natural polymer hydro-
gels [224]. A potential explanation for the latest rise in interest in synthetic polymer hy-
drogels is the simplicity of their industrial manufacture and their ability to be highly mod-
ified, allowing multiple geometries for the construction of tissues [224].
• Poly(ethylene Glycol) (PEG)
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PEG consists of ethylene oxide monomers in its simplest version. PEG is an extremely
diverse synthetic substance that is popular in the biomedical field due to its ease of cus-
tomization [258]. Different degrees of polymerization and varying molecular weights can
considerably modify the mechanical characteristics of PEG. The polymer can also have
different names depending on its molecular weight, such as PEG with a Mw less than 20
kDa, poly(ethylene oxide) (PEO) a Mw over 20 kDa or poly(oxyethylene) for any Mw
[258]. As a non-viscous preliminary solution, PEG is an appealing starting material from
which to produce cell inks, because it can be especially adapted for tissue engineering
purposes. Photopolymerization is the most common technique to produce PEG hydro-
gels, in which light is employed to transfer liquid PEG macromer mixtures into solid hy-
drogels [306]. PEG acrylates, such as PEG diacrylate (PEGDA), PEG dimethacrylate
(PEGDMA) and multi-armed PEG (n-PEG) acrylate (n-PEG-Acr) are commonly used for
photopolymerization [306]. The utilization of tetrahedral PEG tetracrylates (TetraPACs)
could be applied in an extrusion-based bioprinting procedure [310]. Thiolated hyaluronic
acid linked with TetraPAC, a PEG derivative and agarose microfilaments have been uti-
lized to create hollow vascular conduits by bioprinting. Fibroblast cells from mice (NIH
3T3) have been embedded in this hydrogel blend and exhibited high viability [310]. Pep-
tide-conjugated PEG has been applied to print human mesenchymal stem cells (hMSCs),
where the resulting prints possessed outstanding biocompatibility and the nozzle-based
bioprinter hardly occluded [159]. The inclusion of peptides in PEG has been proven to
enhance cell adhesion and promote several immunomodulatory actions [311]. A cell ink
composed of PEGDA hydrogel and human chondrocytes has been employed for cartilage
repair in a nozzle-based bioprinting technique that facilitates concurrent photopolymeri-
zation and printing [154]. This work takes advantage of the proven ability of PEG hydro-
gel to be biocompatible, to be broken down by the body and not to alter the phenotype of
chondrocytes [154]. Most crucially, the PEG hydrogel’s compressive modulus can be ad-
justed to resemble that of human cartilage [312]. The tunability of hydrogels, especially
biodegradable PEG-based synthetic hydrogels, has been investigated [313]. For example,
a polycaprolactone-poly(ethylene glycol)-polycaprolactone mixture (PCL-PEG-PCL) has
been utilized to build a hydrogel with high elasticity and flexibility to facilitate bioprinting
with a visible light deposition curing mechanism to 3D print mouse fibroblasts (3T3) uti-
lizing an extrusion-based printer [313]. Since a low degradation rate of PEG in vivo has
been reported [314,315], the modification of PEG for bioprinting is a highly relevant issue
that has generated encouraging findings.
• Pluronic®
Poloxamers, most frequently referred to by the trade names Pluronic® and Lutrol®,
belong to the category of amphiphilic triblock copolymers, which means polymers with
hydrophilic and hydrophobic regions. Pluronic is heat-sensitive, and its sol-gel transition
temperature range is wide, spanning from 10 °C to 40 °C [316]. Therefore, Pluronic is gen-
erally stable at room temperature and at human body temperature [316]. Because Pluronic
is a synthetic hydrogel, it has a lot of the biological drawbacks of PEG hydrogels, such as
weak cell adhesion and the impossibility of enzymatic breakdown. Nevertheless, a key
benefit of Pluronic is that it has excellent form retention and is thus precise. It provides
structural reinforcement, which means it is also a suitable substitution material. It tends
to become soluble in liquids, so it is frequently inappropriate for prolonged physical in-
teraction with cells. The nanostructuring of Pluronic is an attempt to preserve its structural
characteristics but also to facilitate a long-term cell culture following bioprinting [317]. A
mixture of Pluronic dimethacrylate and non-modified Pluronic has been taken to create
stable gels through UV crosslinking. The non-modified Pluronic is subsequently elimi-
nated from the cross-linked meshwork so that the quantity of Pluronic interfacing with
the cells can be decreased to improve viability. Methacrylated hyaluronic acid (HAMA)
has been incorporated to replace the material removed by elution, which has the benefit
of imparting biological properties to the material. An outstanding cell viability for a
Cells 2024, 13, 1638 24 of 87
Pluronic-based hydrogel has been demonstrated. A high-performance printable, biocom-
patible hydrogel has been introduced for printing of permeable vascular patterns, com-
posed of Pluronic and GelMA [316]. The more Pluronic that is included in the cell ink, the
more improved the printability is. Pure Pluronic has been utilized as a carrier substrate
for the fabrication of vascular structures. An extrusion bioprinter has been proven to pro-
duce intricate vascular patterns, while cell adhesion and proliferation of HUVECs have
been reported [316].
3.2.3. Hybrid Hydrogels
Hybrid hydrogel networks consist of more than one kind of polymer chain or hydro-
gel mesh that is covalently linked together and can comprise both natural and synthetic
polymers [109]. Hybrid bioprinting is frequently employed to produce increasingly intri-
cate structures and offer increased flexibility in shaping [10,109]. The mixture of PEGDA
with alginate is an exemplary hybrid hydrogel [109]. Whereas the PEGDA structures are
chemically linked, the alginate polymers are ionotropically linked [109]. Even though
these are two separate gelling mechanisms, they combine to create a single structure that
has a higher breaking strength and is more resistant to mechanical loads [109]. Another
commonly encountered example of a hybrid hydrogel system is the polyvinyl alcohol
(PVA)/sodium alginate (SA) hydrogel [199]. The PVA/alginate blend offered enhanced
viscosity and enabled direct 3D printing of rigid scaffolds using a core nozzle tip [199].
4. Organoids in 3D Bioprinting
Organoids are 3D in vitro tissue models derived from stem cells that can accurately
replicate the architecture and functionality of human organs. The ability to generate or-
ganoids that mimic the intricate cellular structure of organs has become an emerging
breakthrough technique in biomedical science and the development of pharmaceuticals.
Conventional methods of organoid cultivation are, however, time-consuming and fre-
quently provide only small amounts of cells, which has resulted in the emergence of the
3D bioprinting of organoids from bioinks that contain suspended cells and intended scaf-
folds. The aim of this section is to give a brief description of the traditional production of
organoids and to discuss their advantages and limitations. It will also provide an over-
view of the current status of the 3D bioprinting of organoids and its possible applications
in the fields of tissue engineering, pharmaceutical screening and regenerative medicine.
4.1. Introduction to the Traditional Culture of Organoids
Organoids represent simple tissue engineered cell-based in vitro culture model sys-
tems that mimic multiple features of the intricate structure and functionality of the respec-
tive in vivo tissue. They can be dissected and examined for basic mechanistic investiga-
tions of development, regeneration and repair in human tissues and can also be applied
in the fields of diagnostics, modeling of diseases, pharmaceutical development and per-
sonalized medicine. Organoids can be derived either from pluripotent or tissue-resident
stem cells, either embryonic or adult, or from progenitor or differentiated cells from
healthy or diseased tissues like tumors.
Stem cells are crucial for sustaining organ size, structure and functionality due to cell
renewal, migration, differentiation and apoptosis [318]. Stem cells are placed in a certain
microenvironment, commonly known as the stem cell niche, to govern the fate of stem
cells [319]. Considering the relevance of these environmental factors, there have been mul-
tiple efforts in tissue engineering to engineer the stem cell niche in vitro to provide high
spatial and temporal support for cell–cell and cell–matrix interfaces and to replicate the
mechanochemical drivers using engineered hydrogels and microdevices [320,321].
As Matrigel, a basement membrane ECM comprising a unique mixture of ECM com-
pounds and growth factors, has been extracted from mouse sarcoma tumors, it has ad-
vanced cell culture systems and has been widely used to support in vitro cell culture [322].
Cells 2024, 13, 1638 25 of 87
It has subsequently been found that Matrigel enables mammary epithelial cells to grow in
three dimensions and create lumens that secrete milk protein [323]. Adult intestinal stem
cells incorporated in Matrigel and containing a tissue-specific cocktail of growth factors
have also been capable of self-organizing into 3D crypt-villus architectures [324].
An organoid consists of a self-organized 3D tissue that is usually derived from (plu-
ripotent, fetal or adult) stem cells and imitates the essential functional, structural and bi-
ological intricacy of an organ [325–327]. The cells that make up the organoids can be
sourced from induced pluripotent stem cells (iPSCs) or tissue-derived cells (TDCs), com-
prising normal stem/progenitor cells, differentiated cells and cancer cells [328]. In com-
parison to traditional 2D cultures and animal models, organoid cultures allow a patient-
specific design of the model and simultaneously replicate in vivo tissue-like architectures
and functionalities in vitro. Organoid cultivation is more easily amenable to tampering
and in-depth biological investigations [329] compared to animal models. Organoid cul-
tures have been utilized for a multitude of applications, most notably in pharmaceutical
research [29,330], personalized concomitant diagnostics [330] and cell therapy [329].
Organoid cultures displaying considerable heterogeneity and varying degrees of
compositional intricacy may suffer from insufficiently guided morphogenesis in the self-
assembly process and are frequently devoid of stromal, vascular and immunological ele-
ments [321,328]. Therefore, there is a strong demand to advance organoid culture by ex-
ploiting the knowledge of organogenesis and the interplay of cells with their cellular and
physical surroundings in the shape of the stem cell niche. Based on this knowledge, bio-
engineering approaches could be devised to accurately guide stem cell choices throughout
organoid development. It is known from investigations into early embryogenesis, for in-
stance, that morphogen gradients control the patterning and development of tissues
[331,332]. With the help of microfluidic devices, the desired concentration gradients can
be generated by diffusion of morphogens, resulting in the targeted cell types with spatial
structuring [331]. In addition to biochemical signals, stem cells also perceive active and
passive forces stemming from their external microenvironment and translate these phys-
ical cues into biochemical reactions [333]. These physical properties result from the matrix,
external forces and/or cell–cell interactions. Instead of depending on a natural or biologi-
cally derived ECM like Matrigel, whose stiffness can only be adjusted to a limited extent,
synthetic hydrogels or other ECM mixtures can be used to manipulate the physical char-
acteristics of the matrix. The friction of the fluid against the plasma membrane can also
apply shear stress to the cells [334]. The dynamic biofluidic surroundings have different
consequences for various cell types according to their extent, direction and frequency
[334]. Microfluidic systems and bioreactors can therefore be used for perfusion on a micro-
and macro-scale [335–337]. Cells are recognized to engage with their neighbors and react
in a collective way toward external signals [338]; topographical signals, like the curvature
and shape of neighboring cells, can influence stem cell decision making [339]. A newly
developed neural tube model has effectively dismantled the folding process and shown
that geometric restraints can drive the ultimate morphology of neural tube-like structures
through micropatterning [340].
It is controversial whether artificially produced cell-based in vitro models like organ-
oids must accurately reproduce the structures and functions of the original in vivo organ.
There is a trend towards reproducing the architecture and functionality of in vivo tissues
in vitro as far as possible to prove the physiological validity of increasingly complex mod-
els. For bioengineers, the artificially generated in vitro models only have to reproduce
certain characteristics of the in vivo tissue that are of particular relevance to physiological
or pathological functionality. It is optimistic to build highly intricate models and antici-
pate that they will precisely replicate the organ of origin in vivo. For the majority of sci-
entific issues, simpler models, such as a model with one or two cells within a monolayer
or 3D culture, are more reliable for mechanistic investigations and applications [341–343]
than more complicated models like assembloids or other multicellular models. Experi-
mental aspects of the structure of organoid-based cultivations, which are divided into four
Cells 2024, 13, 1638 26 of 87
main elements, such as cells, soluble factors, matrix and physical cues, and the discussion
of approaches for integrating these elements are shown in (Figure 5). A discussion of key
considerations for creating more intricate yet resilient organoids, such as cell isolation and
seeding, matrix and soluble factor choices, physical cues and integration has emphasized
the 3D bioprinting process.
Figure 5. Elements of organoid cultivation technique. Setting up an organoid-based culture neces-
sitates deliberations on the key elements that constitute organoid cultures, comprising cells, soluble
factors and matrix, physical cues and the effective incorporation of these elements. ASCs = adult
stem cells; CSCs = cancer stem cells; ECM = extracellular matrix; FGF = fibroblast growth factor;
iPSCs = induced pluripotent stem cells; OoC = organ-on-a-chip; TDCs = tissue-derived cells.
Most of the collective behavior arises with the correct 3D tissue organization and cell
constitution, both of which can be delivered through 3D organoid cultures. In 3D organ-
oid cultivations containing tissue-specific morphogens and growth factors, stem cells, in-
cluding embryonic stem cells, those induced pluripotent stem cells and tissue-specific
adult stem cells, those incorporated in Matrigel or under other experimental settings, per-
form tissue-specific differentiation and morphogenesis and progress to organ-specific tis-
sues. The organoids possess a similar cell constitution, tissue morphology and tissue func-
tionality like their in vivo equivalents; for reviews of advances in organoid systems, see
[46,344,345]. It has been demonstrated that various collective cell behavior patterns have
been reproduced within organoid cultures [346,347]. The process of developing human
organ systems such as the nervous system, the lung system or disease systems, such as
tumor organoids, is based on spatially and temporally controlled interactions of cells de-
rived from different lineages [348]. These interactions take place at an early stage of ges-
tation and are thus not amenable to investigations examining neurodevelopmental phe-
nomena or assessing the effectiveness of drugs which target tissues in their native
Organoid
Soluble factors
•Growth factors and small molecules
•ASCs: Wnt/EGF/HGF,IGF, FGF,
BMP/TGF/ROCK/MAPK
•IPSCs: Activin-A,
BMP4/Wnt/FGF(VEGF/BMP/TGF)
Cell Source
•Tissue-based organoids are generated from
tissue and organ biopsy specimen of
humans and animals
•iPSC-derived organoids are harvested from
in vitro grown cell lines
iPSCs CSCs
TDCs
Physical cues
•Offer ECM support and signal
transduction events
•Replicate the nutrient and waste
diffusion processes of the basement
membrane
•Dynamically adaptable surroundings
similar to stem cell ncihes
Matrix scaffold
•Matrigel, collagen, synthetic ploymeric
hydrogel
•Gels made of ECM compounds
•Synthetic gels offer decoupling of stiffness
and structure and thus enhance variability
Integrating cues
•Implementation of major physiological and
structural cues of the organ
•Bioprinting offers direct placment of
organoids to generate tissues
•OoC and multi-OoC systems
Cells 2024, 13, 1638 27 of 87
environment. Human neural organoids, stem cell-derived 3D cultures that self-organize
and display tissue-like cytoarchitecture and physiology, have been shown to accurately
mimic aspects of brain development in vitro [349–352]. Thus, they are emerging model
systems to provide mechanistic understanding of disease etiology [353,354]. Several neu-
ral organoids have been merged into single integrated tissues, termed neural assembloids,
to enable cell–cell interactions and circuit generation in the developing brain to be mod-
eled [355–361]. Traditionally, the fusion of neuronal organoids is accomplished by manu-
ally placing organoids using a large-diameter pipette tip into a microcentrifuge tube filled
with culture medium, where the individual organoids fuse over the period of several days
to create an assembloid [362]. The construction of these structures provides temporal con-
trollability of the interfaces between the organoids, but the multidimensional spatial con-
trollability of their merging continues to be a huge task. The incorporation of various cell
types into organoids is of great importance not only for the recapitulation of neurodevel-
opmental mechanisms and the investigation of the etiology of neuropsychiatric diseases.
For instance, organoid-based cancer models have proven to be a versatile framework for
preserving inter- and intratumoral heterogeneity, allowing ex vivo examination of pa-
tient-specific tumor propagation [16,363]. Thus far, two main strategies have been estab-
lished to reconstruct the cellular microenvironment of tumor and host in vitro. The first is
through the use of genetic engineering approaches to achieve the induction of oncogenic
mutations and the second is through the co-cultivation of cancer cells with organoid mod-
els of the original tissue or the tissue of the metastasis. These models allow temporal sup-
port of tumor–host tissue interfaces but provide restricted spatial guidance of juxtacrine
and paracrine signaling inside the tumor microenvironment.
4.2. Benefits and Limitations of Organoid Cultures
The advantages of organoid cultures are that they enable high-resolution image-
based assessment of the spatio-temporal dynamics of cell–cell interactions inside the tis-
sue under examination. The high number of specimens that can be easily generated with
organoid cultures allows quantification at the global/atomic scale and systematic investi-
gation of critical stages that collectively modify tissue characteristics. Relative to 2D cell
cultures or animal models, organoids offer a more precise depiction of human tissue and
provide more robust and effective drug screening and functional evaluation using patient-
derived lung organoids [364]. This property is especially useful in the field of cancer re-
search, where organoids can replicate the tumor microenvironment and offer valuable
perspectives on tumor-immune interferences and host–pathogen dynamics utilizing pan-
creatic cancer organoids [365]. Organoids are considered more clinically robust than tra-
ditional models because they can mimic the intricate biological processes of human organs
in vitro. This property facilitates fast functional evaluation of pharmaceuticals and im-
proves the effectiveness of the route from drug identification to clinical implementation
[366]. In addition, organoids constitute a novel stage for in vitro gene editing treatments.
Through the use of CRISPR-Cas9 (either to eliminate a gene or rectify a disease-causing
mutation) and other gene-editing approaches, organoids can be utilized by scientists to
model genetic diseases and evaluate therapeutic options, thereby substantially advancing
the domain of personalized medicine [367,368]. In addition, organoids are suitable for the
investigation of organ evolution and pathophysiology in vitro due to their capacity for
self-renewal and their amenability to genetic manipulation.
The limitations of organoid cultures are that organoids are usually devoid of a vas-
cular system, as it is not possible to implement a circulation system with a flow conditions.
Thus, the organoids are grown under static conditions and most of them lack a vascular
system. Thus, in the center of the organoid are hypoxic conditions When the organoids
are grown over a period of several weeks, the cells inside the nucleus can no longer be
adequately supplied with nutrients and the replacement of waste products is hampered.
Thus, necrosis can occur, which can even lead to changes in the mechanical characteristics
of the organoid and in the cytoarchitecture of cells of the organoid. For these reasons, the
Cells 2024, 13, 1638 28 of 87
broad utility of organoids is restricted. In addition to these obvious constraints, there are
others that only become apparent after further analysis of the organoids, including the
lack of highly pure cell types, restricted maturation, atypical physiology, possibly circuit
formation and the absence of arealization, all of which are characteristics that may com-
promise their usefulness for specific purposes. Organoids show an elevated expression of
cellular stress indicator genes that point to metabolic stress, endoplasmic reticulum
stress/unfolded protein reactivity and electron transport disturbances [369–372]. These
disturbances can lead to alterations in the biochemical and mechanical features of organ-
oid cultures.
Thus far, numerous organoid engineering mechanisms have been described to pro-
mote organoid cultivation and growth, proliferation, differentiation and maturation. The
multiple impacts of factors that operate in the in vivo environment pose a difficulty for
the investigation of causality in animal models. As an alternative model to surmount this
problem, in vitro 3D organoid cultures have emerged that offer a reductionist model and
nevertheless show similarities to in vivo tissue in terms of cellular make-up and tissue
organization. In addition, the combination of 3D organoid cultures with other biological
and mechanobiological techniques enables a complex multi-purpose application. Some
limitations can by optimized as outlined in the following.
4.3. Limitations and Optimizations of Organoid Cultures
A major problem is still the temporal and spatial control of the organoids, such as
cell–cell interactions. The reproducibility, in terms of both morphology and functionality,
of the 3D organoid systems produced continues to be a huge challenge.
Limited maturity and function: While none of the existing organoid model systems
replicate the full physiological program of cell types, maturation, and/or functioning of
the organ in question, they instead feature specific functionality of the tissue they pre-
dominantly make up. The overwhelming majority of tissue-derived organoid models lack
tissue-specific cell types, comprising niche-specific mesenchyme, immune cells, vascula-
ture, innervation or microbiome. Co-cultures of ductal cells and liver mesenchymal cells
have recently been found to reconstruct a section of the architecture of the hepatic portal
vein [373]. A particular difficulty is that not all cell types share the identical proliferation
rates, growth factor demands or even oxygen exposure needs, such as the hypoxia for the
vascular system. Organoids derived from pluripotent stem cells are far more capable of
reconstructing the various cell types and cellular interactions of the evolving organ but
lack the structures and functionality of adult tissue and the maturation of cells. A strategy
that can provide assistance is in vivo transplantation [374]. However, this comes at the
cost of maintaining control over the tissue constructs created. In the meantime, differenti-
ation protocols are being improved to enhance maturation and increase the specific func-
tionality of concern. An additional factor influencing maturation and functionality is the
(in)accessibility of nutrients and the accumulation of dead cells in cavities. This is espe-
cially relevant for iPSC-derived organoids. As the size of the organoids increases, the nu-
trient support of the cells in the core of the organoid is limited, which causes cell death.
This is often the case with organoids that build a denser structure, like brain organoids.
In organoids from tissue that build a hollow cyst, such as cholangiocytes and the pancreas,
dead cells build up in the lumen over time, which is inevitable but can be remedied by the
mechanical fragmentation of the organoids. The continuous fragmentation of the gener-
ated structures hinders the conduct of long-term experiments. Organoids derived from
pluripotent stem cells, in contrast, cannot be fragmented and passaged, and new strategies
are being explored to solve the issue of nutrient accessibility, including long-term preser-
vation of brain slices in vitro [375].
Restricted regulatory influence on heterogeneity: As soon as the cells constitute an
organoid, there is a minimal influence on the behavior of the cells inside the organoid. The
outcome, though in the same experimental setups, is frequently a plethora of phenotypic
features, such as shape, size and cell composition, and not a stereotypic culture. The
Cells 2024, 13, 1638 29 of 87
improvement of morphogenic gradients, tissue-specific cell–ECM interfaces and local bi-
ochemical and biophysical characteristics are indispensable for reducing batch-to-batch
heterogeneity [376]. In the field of organoids, efforts have been made to produce more
elaborate multicellular mature and functional structures by generating assembloids, as is
the case with human cortico-motor assembloids [359]. This type of effort enables the gen-
eration of more complicated structures that combine multiple tissue types with a well-
defined interface, such as the interconnection of cerebral cortex, spine and skeletal muscle
with neuro-muscular junctions, albeit at the cost of reproducibility. As discussed in an-
other recent report on the organoids of the liver, bile ducts and pancreas [377], there is a
reduction in reproducibility in multicellular and cross-tissue organoid systems, as it is
difficult to orchestrate the proliferation and differentiation of multiple cell types. The re-
stricted degree of control of heterogeneity within organoids is disadvantageous for high-
throughput screening approaches and hampers investigations that need imaging with
high spatial and temporal resolution. Rather than building more intricate organoid sys-
tems, simpler models with smaller dimensions are progressively being utilized to recreate
the key tissue structures and functionalities of concern. Versions of ECM mixtures, micro-
structured 2D monocultures or co-cultures [378,379], cell sheets [380], 3D stacked textures
[340] and micro-positioned ECM supports [381,382] facilitate the generation of reproduc-
ible tissue architectures and functionalities with a high level of spatio-temporal control;
for instance, through stretching [383] and osmotic forces [384].
Optimization of ECM formulation: Engineering approaches have been developed
to overcome these constraints. Two main ways to address the need to use non-specific
ECM like Matrigel are the application of synthetic matrices with more full control of both
composition and stiffness, and the use of decellularized tissue to produce tissue-specific
matrices [385,386]. Significant progress is being made to establish chemically defined,
GMP-compatible ECMs that permit the growth and long-term propagation of human or-
ganoids. In this respect, some progress has been made with human pancreatic, intestinal
and colon cancer organoids, which were able to proliferate in a fully defined dextran-
based ECM, but failed to grow long-term [320,387].
In the following subsection future frontiers are discussed and an outlook is pre-
sented. There is a tendency to create more advanced models that mimic structure and
functionality in vivo as closely as achievable in terms of cell types that undergo reconsti-
tution over time, the architecture of the tissue, quantifiable molecular processes and phe-
notypic functionality. Instead of concentrating solely on the most relevant landmarks or
functional testing, an architectural comparison with native tissue is also required. Using
the hepatocyte organoids as an example [388], the functions of the hepatocytes are re-
tained, but the architecture of the liver tissue does not correspond to the native tissue, in
which the hepatocytes are organized in strands. Similarly, organoids like pancreatic or
colon cancer organoids exhibit isotropic growth and develop a cyst instead of the tubular
structure they would otherwise develop in their original tissue. To derive more advanced
functions, organoids with multicellular and cross-tissue structures will be relevant, par-
ticularly in the investigation of cell–cell interactions [389]. In this sense, assemblies and
organs-on-chips are also growing in complexity and are being used more and more
widely.
In contrast, the engineer’s attempt [342,390] was to adopt a simpler reductionist mod-
els defined by the minimal functional modules controlling a complex cell or tissue func-
tion of interest to examine mechano-biological causality in the development or repair, or
to design a rugged system for high-throughput compound screening. The fundamental
assumption is that a complex biological operation is carried out by the orchestrated func-
tioning of a limited set of functional units, each of which is characterized by a small num-
ber of molecules, and by chemical reactions that cause alterations in the physical charac-
teristics of mesoscale, such as subcellular or intercellular tissue/multicellular, structures
linked to the functionalities of concern in the distinct spatio-temporal stage/phase/step.
The bile canaliculi in the liver, for instance, undergo hourly expansion and contraction
Cells 2024, 13, 1638 30 of 87
loops. To investigate the underlying contraction events at high resolution, only areas of
neighboring hepatocytes that constitute the bile canaliculi are examined directly in the
relationship with the overall regulatory mechanism of the adjacent hepatocytes [341,391].
A much larger structure can be built with cholangiocytes than that driven by the minimal
functional modules, but the model will be noisier and more expensive. Every functional
unit is linked to a different one and can be analyzed jointly or separately on various length
scales. Fundamental reductionist models have proven valuable in providing high-resolu-
tion mechanistic insight into morphogenetic processes in tissues, such as in the develop-
ment of defects [343,382,392].
Geometrically limiting the size of the initial 2D seeding template and 3D formation
through micropatterning and promoting 3D cell growth with Matrigel facilitated the in-
ducement of tissue-like neural tube morphogenesis and the generation of highly repro-
ducible neural tubes. This also enabled the identification of the mechanisms of neural tube
convolution and the subsequent modeling of neural tube faults [340]. In another case,
symmetry breaking in a uniform cell sphere and the formation of a Paneth cell is a seminal
step in the early phase of intestinal organoid development. The mechanism has just been
clarified: symmetry breaking is induced via transient activation of the mechanotransducer
YAP1, which triggers lateral suppression of NOTCH-DLL1 signaling [393]. YAP1 activa-
tion have since been precisely regulated by utilizing geometric restraints in hydrogel scaf-
folds to generate uniform and reproducible intestinal microtissues [320].
Organoids can be confined by shrinking the third dimension in a 2.5D culture. The
2.5D culture minimizes the depth-related fluctuations of a typical organoid: diffusion re-
strictions in the hypoxic center, restricted access for medication/transfection agents and
restricted transparency in imaging [394]. Typical 3D limitations are the culture of cells on
curved or patterned surfaces, a flattened or restricted cellular construction [394] and plac-
ing the ECM on a flat cell monolayer at a high confluence, which would drag the cells
upwards and force increased cell–cell interactions to attain a 3D cell morphology. Hepato-
cytes within a collagen sandwich have enough contact area to acquire polarity and create
a bile canalicular lumen that constricts in the exact same periodic cycles as it does in vivo,
in the absence of the 3D network, it is wider and cholestatic in comparison to native tissue.
This cell-based model provides a high-resolution breakdown of the bile canalicular con-
traction mechanism into individual steps and an insight into the molecular mechanism
that governs phase transitions [341,395]. Similarly, there could be more artificial organoid
models using CRISPR-edited cells to model diseases, even though these cells and models
remain synthetic. In addition to technological progress in generating more physiologically
valid, rugged and simple-to-use organoid models; however, the impact on applications is
expected to be larger. While there have been discussions about substituting animal testing,
these efforts have not yet resulted in specific interventions. Organoids capable of repro-
ducing the complex physiological processes in vivo have also increased trust that the new
alternative approaches are now feasible choices. Results from animal research are increas-
ingly being transferred to human organoids to gain a deeper insight into human biology
and pathophysiology. Consequently, organoids could be used on a large scale as cell re-
sources for cell therapies, regenerative medicine, in-vitro diagnostics and pharmaceutical
research.
4.4. 3D Bioprinting for Organoid Generation
Collective tissue behaviors, spanning from morphogenesis to the infiltration of can-
cers, rely on the interactions between cells and cells and between cells and their microen-
vironment [396]. These processes are gradually being mapped in self-organizing organoid
and assembloid models [397]. Biofabrication of 3D tissues that replicate organ-specific ar-
chitecture and functionality would benefit from temporal and spatial support of cell–cell
interactions. While bioprinting is theoretically able to deliver this level of control, it is not
well suited to organoids with conserved cytoarchitecture, which are prone to plastic de-
formation. A platform named spatially patterned organoid transfer (SPOT) has been
Cells 2024, 13, 1638 31 of 87
created, which comprises a hydrogel loaded with iron oxide nanoparticles and a magnet-
ized 3D printer and facilitates the regulated lifting, transportation and placement of or-
ganoids (Figure 6) [398]. Cellulose nanofibers are identified both as an optimal biomaterial
for wrapping organoids with magnetic nanoparticles and as a shear-thinning, self-healing
carrier hydrogel to sustain spatial placement of organoids to ease assembloid formation.
Figure 6. Schematic drawing of (A) LIFT and (B) AAB bioprinting techniques for spheroids. Alg =
alignate
SPOT is used to generate accurately arranged assembloids consisting of neural or-
ganoids from human pluripotent stem cells and glioma organoids from patients. In this
way, the potential of the SPOT platform to engineer assembloids that can reconstruct im-
portant developmental processes and causes of disease has been showcased. Three-di-
mensional bioprinting, a technique in which cells, frequently with supporting biomateri-
als, are laid down and assembled into tissues, has been used to achieve control over the
spatial organization of spheroids and organoids. The earlier versions of spheroid bioprint-
ing demonstrated the layer-by-layer extrusion of cell aggregates or cylindrical rods
[310,399–402]. These pioneering efforts used primary cell spheroids that had no internal
cytoarchitecture, were generally limited to less than 500 μm in diameter and were antici-
pated to have standardized sizes so that clogging of the nozzles could be avoided [403].
Organ building blocks (OBBs) printing has since been divided into two different types of
approaches (Table 1) [404]: The first approach is continuous bioprinting, in which the
OBBs are enclosed in the bioink or the supporting scaffold (Figure 6) [405,406], and the
second approach is aspiration-assisted bioprinting (AAB) (Figure 6), in which individual
OBBs are manipulated through vacuum pressure (Table 1) [88,407]. While continuous bi-
oprinting of neural organoids can generate thick, patterned tissue architectures [406], it is
constrained by the inability to accommodate the placement of individual OBBs and the
high expense incurred in generating sufficient numbers of OBBs to colonize the bioink or
scaffold. Although the processing throughput is considerably reduced, AAB might be
more appropriate for spatially structuring the merging of neuronal assembloids in 3D, as
it is able to govern the precise 3D location of every OBB. Nevertheless, it was found that
AAB is unsuitable for the production of neuronal assembloids because neuronal organ-
oids possess large diameters, a comparatively weak surface tension and a tendency to
AAB
Needle
Spheroid
reservoir
Application
of
aspiration
ofabout 70
mmHg
Lifting a
spheroid
A
B
Gel
compart-
ment
Alg
micro-
gels
Transfer to
Algmicrogels
Remove
aspiration
Spheroid
placed
withingel
Bioprinting of
spheroidin
Algmicrogels
Lifting
needle
Bioprinted
spheroid
LIFT
Magnetic
ink
Magnetic
particles
Cellulose
nanofibers
Ink
extrusion
Scaffold
support
Magnetic path
Magnetic rod
Magnet on
Electromagnet
Magnet off
Magnet on
Spheroid
Cells 2024, 13, 1638 32 of 87
plastic deformation and breakdown at quite low vacuum forces [408]. An approach
termed spatially patterned organoid transfer (SPOT) eases the engineering of neuronal
assembloids in 3D with precise spatial controllability across OBB fusion. SPOT uses a cel-
lulose nanofiber (CNF) hydrogel loaded with magnetic nanoparticles (MNP), a CNF sup-
port scaffold enclosed in a tailored container and a magnetized 3D printer to guide the
spatial placement of the OBBs (Table 1). After merging, the generated assembloid can be
detached from the carrier by bioorthogonal, demand-driven disassembly of the CNF scaf-
fold. SPOT is utilized to constrain the spatial positioning of OBBs in two classes of neu-
ronal assembloids. First, for assembloids utilized in neurodevelopmental phenomenon
trials, SPOT is exploited to assist in the construction of assembloids consisting of dorsal
and ventral forebrain organoids that enable in vitro interneuron migration and integration
assays in the cortex. Second, SPOT is applied to assembloids used in translational research
on disease progression and therapeutic effectiveness to generate tissues that integrate or-
ganoids from human brain tumors into neural organoids. Overall, SPOT has the capability
to accurately and reproducibly guide the spatial dynamics of assembloid assembly and
thus provide a high-performance framework for building intricate in vitro models of the
human brain.
Nevertheless, the synergy of advancements in OBB generation with innovations in
biofabrication will be key moving forward as increasing complex interactions between
multiple lineages are to be replicated in vitro [409]. A bioprinting system for organoids,
such a SPOT, has been designed to place individual OBBs in 3D space while maintaining
both a high level of spatial containment and the internal cytoarchitecture. The placement
of these OBBs is accomplished through the utilization of an MNP-loaded, bioinert hydro-
gel that encapsulates the targeted tissue and enables electromagnetically facilitated uplift,
transfer and placement within a hydrogel supporting scaffold. OBBs can merge and build
assembloids inside this matrix. SPOT can be applied to design neuronal assembloids that
function as in vitro models both for a phenomenon of neurodevelopment, that is the mi-
gration and integration of interneurons into the pallium, and for the advancement of neu-
ronal diseases, that is the invasion of cancer cells into various brain areas. This magnetic
bioprinting technique is based on existing pick-and-place biofabrication techniques [410],
such as those employed in AAB [88,407]. In comparison to vacuum aspiration-managed
OBB printing, SPOT decreases concentrated force positioning on the tissue surface and is
therefore ideal for OBBs with low deformation resistance and for use in applications
where the cytoarchitecture of the OBB is important for physiology. While AAB relies on
the manual picking of OBBs within a reservoir of media, SPOT uses a customized chip
layout with microwells for each OBB. This enables the usage of G-code to automate the
finding, picking up and placing of the OBBs at a certain location in the support pool. It is
worth noting that the merging of OBBs has also been achieved previously using the Ken-
zan method, whereby an OBB is aspirated, spiked with a metal microneedle and merged
with other OBBs over the course of several such punctures to form a single assembloid
[411]. The OBB approach has been automated and marketed, but its dependence on per-
foration of the OBB and the associated deformation of the perforated OBB makes it very
difficult to apply to OBBs with preserved, biologically relevant cytoarchitecture. In addi-
tion, the complexity of OBB geometries that can be generated with this approach is re-
stricted due to the stiffness of the needles. Overall, SPOT is a key improvement over other
OBB printing technologies as it enables spatial accuracy in 3D without harming the con-
stitutive OBBs.
It has already been demonstrated that magnetic forces can facilitate the creation of
patterned 3D tissue from individual cells in a procedure referred to as magnetic levitation,
which has meanwhile been marketed [412,413]. Although both magnetic levitation and
SPOT are based on MNPs, there are a number of fundamental distinctions between the
two platforms. First, magnetic levitation maneuvers individual cells into a chosen geom-
etry. SPOT enables the guided motion of whole spheroids or organoids and is thus partic-
ularly suitable for applications where the cytoarchitecture of an OBB is key to its model
Cells 2024, 13, 1638 33 of 87
accuracy. Secondly, magnetic levitation is predicated on the cellular uptake of a bioinor-
ganic hydrogel containing iron oxide, whereas SPOT temporarily coats the surface of an
OBB with an MNP-laden hydrogel. This transient MNP exposure restricts the ability of
OBBs to experience MNP-driven alterations in cellular phenotype. Hence, compared to
this former magnetic bioprinting approach, SPOT is especially useful for the production
of assembloids from organoids with retained cellular architectures.
SPOT seeks to offer a complementary approach to conventional assembloid creation
concepts that rely on the merging of organoids as a result of entrapment in a microcentri-
fuge tube [414]. Although these protocols utilize reagents and equipment easily attainable
in the majority of biology laboratories, the ease of assembly itself restricts the amount of
control that can be exerted over the spatial placement of the OBBs. In addition, linear as-
sembloids from up to three separate OBBs have been verified [415], whereas the creation
of assembloids in X, Y and Z dimensions is still a huge challenge. SPOT has the capability
to be an advancement over existing OBB assembly approaches, as the 3D printer modified
with an electromagnet enables the operator to manage the placement of multiple OBBs
across three dimensions. The SPOT platform is designed to be precise, scalable and adapt-
able to specific needs. Several technical steps have been indicated that may be of potential
interest to those seeking to integrate them into their experimental operations. First, the
MNP concentration, deposition time, magnetic rod diameter and magnetic field strength
need to be properly adjusted for the largest OBB within an experiment. Second, while
SPOT can cover a 300–3000 μm span of OBB diameters, it has difficulty precisely isolating
OBBs below 300 μm. Optimizing the connection of these organoids to the magnetic rod
might solve this constraint. Since this bioprinting framework is OBB-agnostic, it can be
used in a broad spectrum of biological systems, wherein signaling from different cell
types, cell lines and oncogenic capacity is important. SPOT is employed to construct mul-
tiregional neuronal assembloids composed of regionalized components of neuronal cir-
cuits and tumor–host assembloids in which the proportion and placement of each OBB
can be manipulated in a controllable manner. The combination of the SPOT platform with
spatially resolved single-cell RNA sequencing, multiplex time-lapse immunofluorescence
and imaging mass cytometry has the potential to provide powerful mechanistic evidence
on the spatio-temporal dynamics of infiltrating tumors.
The tremendous impact and potential benefits of 3D bioprinted organoids are enor-
mous, and as the technology is further developed, more uses in disease modelling, phar-
maceutical research and regenerative medicine will be realized (Table 1). There are several
hurdles that still need to be overcome before 3D-bioprinted organoids can be imple-
mented on a routine basis in the hospital. Nevertheless, the field of organoid 3D bioprint-
ing has an encouraging future and has the capacity to transform the area of tissue engi-
neering and regenerative medicine.
In the field of nanotoxicology, organoid-based scaffolds have been employed for
long-term investigations in immortalized cell lines [408]. The toxicity of nanoparticles in-
gested through physical contact or when inhaled is a major public health issue. It is im-
perative to perform continuous assessment of the toxicity of nanomaterials. In vitro nano-
toxicology investigations are usually restricted to two dimensions Even though 3D bi-
oprinting has recently been used for three-dimensional cultures related to the liberation
of medicines and tissue regeneration, not much is understood about its application for
nanotoxicology testing. Organoid-based scaffolds have thus been established for long-
term studies in immortalized cell lines with the goal of mimicking the exposure of lung
cells toward nanoparticles. Viscous, cell-loaded material is printed using a customized 3D
bioprinter and then irradiated with either fluorescent latex with a diameter of 40 nm or
silver nanoparticles with a diameter of 11 to 14 nm. The administered fluorescent nano-
particles can diffuse in the 3D-printed frameworks, while this has not been the situation
with the unprinted frameworks. A marked increase in cell viability of 3D versus 2D cul-
tures being challenged with silver nanoparticles has been detected. This demonstrates tox-
icological reactions that mimic in vivo experiments, like inhaled silver nanoparticles. The
Cells 2024, 13, 1638 34 of 87
findings provide a new prospect in 3D protocols for nanotoxicology investigations that
avoid animal testing. Toxicological and nanomedical investigations necessitate the oblig-
atory step of in vivo studies [416,417]. This is due to the requirement to completely com-
prehend the bioavailability, fate and biodistribution of nanoparticles throughout and
post-exposure, as well as the local and systemic effects they cause. The foregoing step can
be largely attenuated by the employment of organoids [304], which replicate the intricate
microarchitecture of ECM constituents and the interactions between different cell types
adequately to reproduce biological functionalities [55], thereby decreasing the amount of
animal testing required for toxicological/pharmacological preclinical evaluation [418].
Organoids can be produced using new techniques such as additive manufacturing
[54], which, through the use of a 3D bioprinter [419], introduces an important novelty in
the field of in vitro tissue regeneration and examination [420], by replicating the in vivo
environment in both mono- and multicellular culture [421,422]. Extrusion-based 3D bi-
oprinting (robocasting or direct ink writing technologies) offers the possibility of produc-
ing cell-loaded scaffolds based on biocompatible hydrogels and enables fast, sterile and
reproducible manufacturing processes. [423,424]. Only very few investigations on nano-
toxicological analysis using cell-seeded or cell-loaded scaffolds from the 3D printer have
been conducted, however, perhaps due to the intricacy of reproducing and guiding bi-
oprinted “living” multilayers [425,426]. A tailored and cost-effective 3D bioprinter has
been utilized to evaluate the advantages and limitations of bioprinted cell-loaded hydro-
gel scaffolds [427] or nanotoxicology and nanomedicine investigations or OBST (organ-
oid-based scaffolds for toxicology studies). Hydrogels based on alginate/gelatin/Matrigel
at various concentrations were assessed to pick the one most suitable to maintain cell vi-
ability and print the cell-loaded scaffolds with a conventional honeycomb structure [428].
In the second phase, the viability of the cells has been characterized, and the findings in-
dicated that the cells could grow in the OBST for 21 days without significant operator
interference, proving that their hydrogel composition preserves the cells for a longer pe-
riod of time and decreases lipid peroxidation. Ultimately, the nanoparticles applied to the
OBST have been characterized by a diffuse engagement with the bioprinted cells, which
produced a similar toxicological reaction as the in vivo tests with AgNPs. Atoxic carboxyl-
modified fluorescent nanoparticles have been used for mapping the distribution within
the OBST by two-photon microscopy [429], whereas AgNPs were used due to their known
cytotoxicity [430–432]. The proposed OBST technique offers various advantages for nano-
toxicology/anomedical studies: first, cells can survive for a longer time without undergo-
ing passages; second, nanoparticles can disperse and diffuse in the cell-loaded multilayer
by imitating in vivo exposure; third, nanoparticles arrive at the 3D-printed cells in all lay-
ers with a significant increase in internalization time in comparison to the non-printed
and conventional 2D cultures and fourth, there is a different dose/reaction of 3D-printed
cell multilayer toward silver nanoparticles (AgNPs) compared to 2D, resembling in vivo
data in zebrafish [430], in insects [433] and rodents [434,435].
To evaluate nanotoxicology studies using 3D-printed cell-laden scaffolds that repli-
cate real cell–cell mixing in vivo [139] and ECM production [436,437], which is difficult to
see in 2D, a 3D bioprinter has been created and constructed. The 3D bioprinter has been
conceptualized and built to reduce the chance of contamination while extruding the hy-
drogel with minimal extrusion force and velocity [269]. The 2D expanded cell lines along
with their specific supplemental media have been integrated into the alginate/gelatin-
based hydrogel formulation [438], which has also been augmented with Matrigel [79], and
the final CAD drawings have been constructed based on bio-inspired honeycomb inter-
secting layers [428], which determines the trajectory for OBST manufacturing. All printed
cells remained biologically active, survived for 21 days with minimal operator interfer-
ence and were capable of internalizing nanoparticles that were subsequently applied to
the OBST to simulate the engagement of skin and mucosa with engineered nanomaterials.
In addition, the level of thiobarbituric acid reactive substances in Calu-3 dropped mark-
edly by approximately 90% over a 14-day culture [408], indicating that the hydrogel
Cells 2024, 13, 1638 35 of 87
composition can preserve the 3D-printed cells for long-term studies. Thiobarbituric acid
reactive substances pointed out that an adaptation period to the 3D environment has been
necessary for all cell lines under investigation. As previously stated for cell proliferation,
the cells regained their normal biological functionality after a short period of time and
reverted to a condition of decreased membrane lipid peroxidation [408]. The amount of
dead cells stayed minimal throughout this time period. It has also been essential to guar-
antee uniform cell deposition when creating an OBST. Unprinted scaffolds showed not
merely unevenly distributed cells, but also air bubbles and uneven internalization of the
nanoparticles by the OBST. It is finally noteworthy that the 2D in vitro toxicity of high
doses of AgNPs to mammalian cells is unquestioned [439,440], the OBST obtained find-
ings are similar to the in vivo toxicological results in rats following inhalation of AgNP.
In addition, a less pronounced decrease in viability of 3D-printed cells subjected to AgNPs
has been seen, and the ability to study the same cells imbedded in a 3D architecture for
weeks could speed up the process from lab to patient. Cell-loaded 3D-printed multilayers
offer the possibility to investigate lipid peroxidation longitudinally, as the damage caused
by oxidative stress is similar to that of lung tissue in vivo [417] and, and according to the
physicochemical properties of the nanoparticles [441], they can be used for toxicological
or nanomedical investigations, as the nanoparticles can diffuse effectively into the printed
layers. Although the OBST technique is at an early stage and needs further studies, it could
be an effective instrument for nanotoxicology studies where the cells incorporated in the
3D hydrogel are active and can interact with the nanoparticles [442] and the scaffold [443].
In conclusion, the research emphasizes the significant differences between 2D and 3D
data, suggesting that consideration needs to be given to reviewing tactics in the fields of
nanotoxicology and nanomedicine to account for possible impacts on cell morphology
and cell–cell interactions in a 3D environment. Ultimately, the technology can aid the de-
velopment of safer and more powerful nanomedicine and represent a useful tool for sci-
entists in the nanotoxicology field. There are certain limitations, nevertheless. For exam-
ple, the number of cells seeded strictly relies on the 3D printing parameters and is usually
lower than the amount of cells embedded in the hydrogel-loaded syringe; this is due to
the remaining volume in the syringe’s Luer lock and the force applied during extrusion,
which is minimized but mechanically destroys the cells. Consequently, at least two major
questions arise: How can scaffolds be employed in the 3D organoid culture? How is bi-
oprinting helpful?
5. Three-Dimensional Bioprinting of Complex Geometric Models with Tumor Organ-
oids Serving as Structural Elements
Three-dimensional bioprinting is a promising biomanufacturing process that pro-
duces multiscale tissue models of remarkable accuracy and physio-mimetic performance
[444]. The most widespread technique is extrusion-based bioprinting, also referred to as
bioplotting. In this technique, endless filaments of bioink are pushed or mechanically ex-
truded via printer nozzles into predefined 3D structures [445]. Whereas in conventional
bioprinting through extrusion, the bioink is deposited layer by layer onto a carrier sub-
strate. In a new technique referred to as embedded bioprinting [446], the bioink is placed
in a slurry pool in which it is held against gravity, enabling the construction of intricate
textures, such as blood and lymphatic vessels [447], which also extrudes cell-laden bioinks
via nozzles. In this process, however, thermal, electrostatic, or piezoelectric techniques are
used to supply droplets of bioink instead of continuous filaments so as to obtain a higher
level of resolution of the deposit. There are also other nozzle-independent bioprinting
techniques, like acoustic printing [448]. These droplet administration techniques can de-
liver cancer cell-loaded bioink droplets with enhanced spatial resolution into the tumor-
specific CAF-harboring scaffold [448]. The most frequently employed technique is extru-
sion-based multi-material bioprinting, where coaxial printheads or multiple nozzles are
deployed to extrude several bioinks with varying cell and material combinations [449].
Inkjet and acoustic printing are emerging bioprinting techniques that facilitate multi-
Cells 2024, 13, 1638 36 of 87
material engineering to enable the concurrent deposit of various cellular components and
matrix substances [450]. These advances allow improved design of advanced tumor mod-
els with more heterogeneous compounds to mimic the fundamental cell–cell matrix inter-
play.
A miniature brain model has been designed with the aid of the multi-nozzle bioprint-
ing technique in which GBM cells and macrophages have been placed in specific compart-
ments of the GelMA [451]. This model has consistently replicated the cellular interplay
between neoplastic and immune cells, involving the enrollment and the transformation of
macrophages into the GBM-associated macrophage phenotype and the invasiveness of
GBM cells into the mini-brain tissue, correlating well with clinical transcriptome results.
Most investigations are performed with monodispersed cancer cells that serve as bioprint-
ing building elements. Monodisperse cells, however, are unable to accurately mimic tu-
mor propagation, as volumetric cancer cells hardly ever occur in isolation [452]. Organoid-
based bioprinting constitutes a viable mechanism to incorporate miniaturized tumor ag-
gregates within a heterogeneous 3D cavity with assisting hydrogels and stromal cells. The
combined effect of these features enables self-organization of tumor-sized anatomy with
hierarchical functional modules [453], which provides a more accurate reflection of intrin-
sic TME features. A growing demand and unparalleled possibilities are emerging for the
creation of novel and more efficient tissue engineering techniques, among which 3D bi-
oprinting is considered one of the most encouraging. Although biomaterial-dependent 3D
bioprinting is progressing continuously, it is still slow to yield the expected therapeutic
outcomes. Alternative “scaffold-free” 3D bioprinting methods are therefore currently un-
dergoing rapid progress. The readiness of bioprinting techniques and the quality of the
bioprinted structures should be assessed before they can be employed for therapeutic fea-
tures [48,454].
5.1. Scaffold-Free Bioprinting of Tumor Spheroids
Various bioprinting techniques for the accurate placement of cell assemblies have
been established. The initial approaches of these techniques employed a more accessible
form of tumor aggregates termed tumor spheroids that are more robust for the bioprinting
procedure compared with PDOs. Among the initial concepts of bioprinting of tumor ag-
gregates is the Kenzan technique. In this technique, the aggregates are positioned on a
microneedle arrangement with the aid of a robotic stage in a cohesive template designed
in advance. The microneedle array acts as a guide for the aggregates to merge into a larger
cellular entity and initiate the synthesis of their individual ECM scaffold. The Kenzan
technique eliminates the reliance on biomaterial frameworks, which is why it is also re-
ferred to as a framework-free method. It is suitable to produce microtissues with elevated
cell density and direct cell–cell interfaces, including adipose tissue, cartilage, nerves and
heart muscle-like assemblies. In tumor sculpting, the Kenzan technique offers a powerful
way to replicate the two-way interaction between tumor mass and adjacent mature tis-
sues. Using this approach, the neuro-like parenchyma enveloping glioma tumor sphe-
roids has been recreated to assess real-time invasion of the cancer [455]. Nevertheless, this
model cannot represent the relevant gliosis features of glioma disease pathology as it is
not able to reproduce various cell–ECM compounds and interferences [455]. The Kenzan
method is restricted in its spatial resolution and does not control for biochemical/biophys-
ical issues associated with the ECM framework. In addition, this technique has not yet
been implemented on tumor organoids, probably since the mechanical interruption of this
procedure can lead to exaggerated injury to the tumor organoids, which would lead to
decreased survivability. Another scaffold-free technique is the liquid-based singulariza-
tion approach, where single tumor aggregates are trapped and liberated one after the
other through a fluid-controlled rear pressure [456]. This singularization unit can be in-
corporated into bioprinting nozzles to accurately place individual spheroids into 3D bi-
oprinting frameworks. This approach facilitates the assembling of tumor organoids with
the other key features of the intricate TME and guarantees high accuracy of printing [456].
Cells 2024, 13, 1638 37 of 87
It also allows the creation of multi-level structures with tissue-specific forms and quanti-
ties. For instance, the liquid-based singularization method has been applied to produce a
macroscale ovarian model in which heterotypic spheroids comprising both ovarian ade-
nocarcinoma cells and fibroblasts have been assembled in a 3D-printed hydrogel frame-
work [457]. The growing scale and sophistication of TME has decreased the sensitivity to
doxorubicin in comparison to individual spheroid units. Nevertheless, the slow operating
pace restricts the practicability of this technique for large-scale and high-performance pro-
cessing.
5.2. Scaffold-Based Bioprinting
In bioprinting processes, hydrogels are often used as carrier substrates in 3D bi-
oprinting. These carrier media consist of so-called sacrificial inks and carrier baths, which
can temporarily mechanically reinforce the bioink during the printing process. Hydrogels
featuring reversible sol-gel phase properties have been employed as temporary sacrificial
inks to enable printing of cavities, incorporating permeable channels [444,445]. Hydrogels
with thixotropic mechanical characteristics may be employed as carrier baths that facili-
tate the printing of intricate structural characteristics with enhanced print detail while
maintaining physical constraint throughout printing [281,458,459]. The utilization of gel-
phase carrier media has significantly expanded both the level of complexity of printed
geometries and the spectrum of materials that could be applied as bioinks [82,460]. Poly(ε-
caprolactone) (PCL) and GelMA are two well-known biomaterials for printing. Multiple
scaffold-based bioprinting efforts have employed biocompatible hydrogels to embed
spheroids/organoids and several stromal cell types. This approach relies on the bioprint-
ing of monodisperse cells by mitigating mechanical disturbances and simultaneously
guaranteeing correct dimensional placement, intricate geometry and hierarchical diver-
sity [407]. Due to their excellent biocompatibility, scaffold-based bioprinting procedures
have shown considerable benefit for the generation of tumor organoids, where also me-
chanical characteristics are important. Bioprinting methods employed for framework-
based bioprinting involve drop-on-demand, acoustic bioprinting and extrusion bioprint-
ing. The models produced with every technique differ in terms of scale. The drop-on-de-
mand process can manufacture smaller-sized models ranging from 100 to 1000 µm, with
a droplet size from 42 to 960 µm [461], while extrusion bioprinting creates larger-sized
models ranging from 10 nm to 100 nm with a nozzle diameter ranging from 260 µm to
1200 µm [462,463]. The drop-on-demand method enables the production of textures with
high output and a high level of uniformity. This method is generally ideal for the printing
of very small structures measuring 100–1000 μm and utilizing low viscosity bioinks. With
this setup, bladder tumor assembloids harboring PDOs originating from luminal/basal
phenotypes and fostering CAFs and ECs could be established [464]. Accurate recapitula-
tion of the TME resulted in the tumor assembloids having similar structural features to
the original cancer tissue with interconnected vasculature [465]. The assembloids also ef-
fectively prevented the transition of phenotype from luminal to basal, which invariably
arises in the long-term culture of organoids of luminal urinary tract carcinomas [466].
Acoustic bioprinting represents a further method for the droplet-based administra-
tion of tumor spheroids/organoids in a physically defined way [467]. In acoustic bioprint-
ing, the droplets are expelled through a mild sound field. This technique can shield living
cells from harmful stress agents like heat, intense pressure, high tension and substantial
shear forces that are present in alternative bioprinting approaches [468,469]. Nevertheless,
acoustic bioprinting is limited in its capacity to expel droplets of a high viscosity bioink,
in a similar way as the drop-on-demand technique [468]. A patient-derived colorectal can-
cer microtissue has been bioprinted with tumor and healthy organoids via droplet release
[467]. Extrusion bioprinting can also be expanded to the bioprinting of spheroids/organ-
oids [470]. In comparison to drop-on-demand and acoustic bioprinting, a broader spec-
trum of bioinks can be utilized in this approach as shear-thinning hydrogels can be em-
ployed. This strategy also allows manipulation of several materials and thus supports the
Cells 2024, 13, 1638 38 of 87
development of a tumor-sized assembly with various ECM compounds and diverse cell
types [471]. A co-culture model of breast cancer preshaped spheroids and ECs has been
fabricated via extrusion bioprinting [472]. This model demonstrated increased tolerance
to paclitaxel therapy when compared to a monodisperse cell printing model, emphasizing
the significance of cell association in the therapeutic outcome [471]. Although advances
have been achieved to this point, the bioprinting of tumor spheroids/organoids needs ad-
ditional fine-tuning to accurately adjust the characteristics of the bioink and broaden the
manufacturing options. Volumetric bioprinting (VBP), where liquid resins are photopol-
ymerized into a volumetric 3D pattern with light, has emerged as a high-performance
bioprinting method with high resolution and manufacturing capability for the fast build-
up of demanding patterns. VBP takes computed tomography as its inspiration and uses a
digital light processing engine to create a sequence of 2D light designs projected onto the
paint reservoir from various angles, causing the build-up of light dose and solidifying the
resin as it approaches the gelling level. Bernal and colleagues [473] used VBP and liver
organoids have been employed to fabricate intricate, centimeter-sized hepatic reconstruc-
tions in seconds. This advanced method is without nozzles or layers and guarantees a
high degree of survivability and form accuracy of the emerging organoids [473]. While
VBP continues to suffer from the issue of light scattering and is therefore not applied for
tumor modeling, it is nevertheless a powerful tool for replicating human-scale cancer
models [473].
5.3. Mechanical Cues of Scaffolds Impact Organoids
Classical strategies for building organoids still cannot accurately reproduce the key
features of real organs as it is challenging to regulate the self-organization of cells in vitro.
The current constraint stems from the impossibility to manipulate the organoid environ-
ment in the classical production of organoids and the scarcity of information and actual
measurements of the biomechanical characteristics of tissues in developing organs, such
as the human brain, or in diseased organs suffering from cancer, injury or inflammation
[474–478]. Nevertheless, the knowledge of mechanical characteristics has increased con-
tinuously due to new developments in biophysical techniques [241,475,476,479–482].
Thus, this information on mechanical cues can be used to mimic them in scaffolds for the
generation of organoids. Mechanical characteristics of 3D bioprinted scaffolds, such as
softness (or rigidity) or curvature, can contribute to the morphology and functionality of
organoids [483,484]. For instance, it has been found that the conversion of a round intes-
tinal stem cell (ISC) colony into a crypt-containing organoid inside a synthetic hydrogel
necessitates a softening of the matrix [320]. Nevertheless, the global matrix softening
model employed led to stochastic and spatially unsupervised budding, just like in tradi-
tional organoid cultures using native ECM 3D matrices [324,485].
6. Organ-on-a-Chip Convergence for Optimized TME and Bioprinting
Organ-on-a-chip technique can be partly bioprinted. The combination of organ-on-a-
chip technique and bioprinting seems to be interesting for developing more intricate
model systems in a reliable manner. Intercellular communications are critical for the
proper operation and evolution of organisms. Cells can communicate with one another
both directly (in physical contact) and indirectly (paracrine signal transmission). These
interactions constitute the fundamental elements of physiological communication and are
vital for the generation of tissue, the immune response, homeostasis and regeneration.
During direct cellular communication, cell surface contact can involve gap junctions, cell
adhesion, tunneling nanotubes and ligand–receptor signaling. In indirect cellular commu-
nication, cellular messages are exchanged by signaling extracellular vesicles, including
exosomes and ectosomes, cytokines, chemokines, growth factors, miRNAs and metabo-
lites. They all contribute to the development of tissue and physiological responses [486].
A healthy immune system can accurately recognize and eliminate cancer progenitor cells
before they cause damage via a mechanism referred to as tumor immunosurveillance.
Cells 2024, 13, 1638 39 of 87
Numerous extrinsic and intrinsic agents interfere with the communication between the
immune system and cancer progenitor cells, leading to tumorigenesis. Cancer cells escape
the host immune system and perturb it, resulting in an immunosuppressive TME. The
TME consists of a complicated ecosystem of cancer cells, immune cells, fibroblasts, ECMs
and regulatory molecules. The immune cell–cancer cell exchange within the TME evolves
and can lead to the generation of either pro- or anti-tumorigenesis [487]. Reestablishing
the immune response and the interaction of healthy cell–cell communication within the
TME is an integral part of cancer immunotherapy [488]. While a variety of other immuno-
therapies such as immune checkpoint inhibitors, oncolytic viruses, bispecific T-cell engag-
ers, cytokine therapies and adoptive cell therapies have been identified, the central idea
of immunotherapy is to reconstitute or reenable the patient’s host anti-tumor immune
system [488]. Despite the potential of these immunotherapies, clinical responsiveness dif-
fers greatly between patients, primarily due to differing combinations of immunosuppres-
sive TMEs and varying interference patterns with cell–cell interactions. Approximately
30–40% of immunotherapy patients achieve a positive response, with only a few obtaining
a permanent response [489]. The TME versatility is mostly linked to the varying degrees
of tumor-infiltrated lymphocytes (TILs) and their functionalities. Certain patients experi-
ence “hot” cancers, where the cancers contain increased TILs, and usually respond favor-
ably to immunotherapy [490]. Conversely, some patients present with “cold” cancers,
with few or virtually no TILs, and these cancers frequently evolve into immunotherapy
resistance [490]. Clarifying the resistance mechanism and determining the aberrant nature
of intercellular communication among cancer cells and immune cells in TMEs is key to
developing more effective immunotherapies. Although advances have been achieved
lately in the bioprinting of organoids/spheroids to implement the build-up of heterogene-
ous TME compounds in a biomimetic matrix [463], there is clearly an enormous amount
of work to be accomplished to replicate intrinsic cancer progression, pharmacokinetics
and pharmacodynamics in vivo. These pathological events rely on tumor-immune com-
munication and interactions among various functional organs, which are greatly stream-
lined in tumor models cultured in a static environment because of the absence of a func-
tional circulatory network [491]. Organ-on-a-chip systems are being investigated as an
advancement in cancer models, as they have excellent capabilities to mimic natural vas-
cular perfusion and blood microcirculation. The organ-on-a-chip comprises complex com-
partments and microchannels capable of being interconnected to mimic multi-organ in-
teraction and facilitate dynamic liquid circulation. This enables the manipulation of the
physical and biochemical variables of TME, such as oxygen levels, pH equilibrium, nutri-
ent distribution, molecular gradients and, most importantly, the flow of circulating cellu-
lar elements [492]. Moreover, organ-on-a-chip technology enables the utilization of novel
integrated physical, biochemical and optical sensor devices. These modular devices can
be utilized for in situ observation of TME variables and dynamic tumor reactions to active
drugs over a longer period of time [493]. In spite of their ability to apply dynamic bio-
chemical and fluidic feedback signals in vitro, conventional methods for fabricating chip
devices, including micromachining and soft lithography, are labor-intensive, time-con-
suming and offer a restricted capacity to accurately dispense living parts [494]. Thus, the
organ-on-a-chip for prototyping bioprinting is a challenging way to tackle these hurdles
[495]. The technology can improve the way it mimics the TME’s intricate characteristics
and support the advancement of powerful cancer treatments [496].
6.1. Assembly Strategies and Shapes for Bioprinting Tumor-on-a-Chip Models
A potential strategy for combining bioprinting and organ-on-a-chip technology uti-
lizes a post-integration technique in which the living parts are bioprinted into pre-engi-
neered microfluidic assemblies. This approach optimizes the application of microfluidic
fabrication methods to create chip carriers with intricate microflow channels and imma-
ture vascular meshes. The downstream bioprinting procedure facilitates the insertion of
biological elements into the chip carrier in a space-time fashion [497]. As part of the post-
Cells 2024, 13, 1638 40 of 87
integration approach, nozzle-based bioprinting techniques like extrusion and inkjet bi-
oprinting are commonly utilized to incorporate multiple bioinks and mixed cell types into
microfluidic systems [498]. Two different types of organ-on-a-chip device, namely Organ-
Trial® Dolores and OrganTrial® Hive, have been presented that employ the post-integra-
tion technique wherein bioprinting models of different types of tissues can be incorpo-
rated into commercially available mounting devices fitted with culture compartments,
microfluidic channels and pulsation-controlled features to enable the dynamic cultivation
of macro-scale tissues and regulate multi-organ interaction [499]. In addition to traditional
chip manufacturing techniques, 3D printing enables the fabrication of chip covers and
microfluidic channels. This technology forms the foundation for a one-step production
approach in which chip components and biological parts can be manufactured at the same
time. The one-step manufacturing approach opens the door for diverse architectural pat-
terns and facilitates the creation of customized tumor-on-a-chip models [500,501], such as
brain tumor models [502]. This strategy, nonetheless, demands multi-material manufac-
turing capabilities that can introduce cell-loaded bioinks and cladding materials into in-
tricate structures concurrently. Extrusion printing meets these demands because it can
employ several print heads for various purposes. For instance, a breakthrough printing
system adopts a photopolymer head, a UV head, a microplasma head and a biologics head
to implement the free-form manufacturing of a tumor-on-a-chip apparatus in a single step
[503]. Stereolithography represents a further established method for the fabrication of
high-resolution chips, although it is extremely challenging to customize for the manufac-
ture of multi-materials. Hybrid printing methods could be more practical and offer more
freedom to create a broader model with a one-step manufacturing approach [504]. An-
other problem is the limited choice of the shell substrate. The circumferential material
utilized in one-step manufacturing needs to be adequately printable, biocompatible and
sustainable to preserve the structural integrity of the complicated chip architecture in
long-term culture environments, which limits the selection of appropriate circumferential
compounds [140]. Compared to the post-integration procedure, the living cells must un-
dergo the lengthy printing process in one-step production, which represents a considera-
ble challenge for the cells’ ability to survive. In addition, the higher degree of difficulty of
the devices leads to lower accessibility, as there are only a few commercial offerings and
most of the devices used in scientific research are manufactured by the scientists them-
selves. Before the one-step procedure can be introduced into clinical practice on a larger
scale, the limitations in material selection, turnaround time and accessibility of the devices
must be eliminated [504]. In the last few years, considerable advances have been achieved
in the creation and utilization of bioprinting-based tumor-on-a-chip models that emulate
TME’s spatially distinct features and dynamic culturing conditions [505]. Several im-
portant engineering aspects play a crucial part in the production process of these intricate
3D cancer models. Due to the advantages in terms of the accessibility of equipment, the
availability of materials and the fast process of cell loading, the post-integration approach
has prevailed in modern 3D bioprinting tumor-on-a-chip models compared to the one-
step option [506]. In addition to the manufacturing process, the selection of cell compo-
nents is another decisive aspect [506]. The use of tumor organoids rather than monodis-
perse individual cells as components is more efficacious in simulating natural tumor–stro-
mal relationships. This design rationale has been accomplished with organoid/spheroid
bioprinting methods that guarantee the construction of 3D tumor assemblies with other
major TME constituents, comprising stromal cells, vascular/lymphatic vasculature and
supportive ECM scaffolds inside the chip units [92,504]. In a 3D bioprinted neuroblas-
toma-on-a-chip model, Ning et al. [504] demonstrated dynamic tumor–vessel interfaces
with success, showing the aggressive response of the tumor and allowing the evaluation
of metabolic, cytokine and gene expression patterns in various TME environments. Not-
withstanding present advances, full recall of the intricate interactions between tumor and
stroma still needs additional work to establish the native TME’s composition and its
unique geometric features. Innovative digital image processing and quantification
Cells 2024, 13, 1638 41 of 87
methods with artificial intelligence are highly beneficial for determining the exact patient-
specific TME criteria [507]. In improving the accuracy and reliability of tumor models by
helping to develop a more informative structural blueprint, these technologies can offer a
powerful tool for investigating cancer biology and drug sensitivity [508].
6.2. Enhancement of Integrative 3D Tumor Models: Vascular System, Immune
Monitoring and Tumor Metastasis
Tumor microvascular structures are key drivers of tumor evolution as they supply
nutrition and oxygen for the survival of the cancer and can create routes for cancer metas-
tasis [492]. Similar to healthy tissue, the development of blood vessels in cancer can occur
through angiogenesis, i.e., the growth of blood vessels through the sprouting or splitting
of already formed blood vessels, and/or through vasculogenesis, i.e., the formation of new
blood vessels from endothelial progenitor cells formed in the bone marrow [509]. The new
neoplastic capillaries, which are connected to established blood vessels, form a hierar-
chical vascular system with vessels of diverse diameters. Mimicking this hierarchical or-
ganization is key to replicating the inherent characteristics of tumor–vessel interfaces and
pharmaceutical kinetics. The organ-on-a-chip approach has enabled two major ap-
proaches to design vascular replicas at various length scales. The first technique is the
vascular patterning, in which prefabricated microfluidic tubules layered with ECs form
microvessels of diameters larger than 100 μm. The second technique involves self-organ-
ization, in which single ECs pass through a process resembling vasculogenesis, allowing
the generation of capillary equivalents with a lumen width of between 15 and 50 μm [492].
Whereas traditional microfluidic techniques are only able to fabricate single-size vessels,
the integration with bioprinting facilitates advances in the direction of integrated hierar-
chical vessels. With the help of multi-nozzle bioprinting, it is possible to introduce substi-
tution material into vessel-like 3D geometries and encase them in EC-loaded hydrogels.
After discarding the voiding material, the void channels can be layered with ECs to create
a sealed vessel-tissue boundary. Three-dimensional bioprinting can be employed to pro-
duce vessel-like 3D embodiments that are far more intricate and lifelike than traditional
“2.5D” designs in which 2D channel patterns are extrapolated into the third axis [510].
Simultaneously, the ECs enclosed in the encapsulating hydrogels are subjected to vascu-
logenesis-like self-organization to create an equivalent capillary mesh. This hierarchical
vascular architecture has been recreated in a neuroblastoma model effectively mimicking
the organization and functionality of cancer tissue with major EC-lined vessels facilitating
perfusion and capillary meshes facilitating tumor–vessel communication [504]. In the en-
gineering of engineered tumor-on-a-chip models, lymphatic vessel systems should also
be addressed. Microcirculatory systems have been recreated in vitro with matched vascu-
lar and lymphatic vessels to improve the simulation of the trafficking kinetics of thera-
peutic cancer drugs [449]. Numerous biomanufacturing techniques such as electrospin-
ning, decellularization of xenogeneic vessels, 3D printing and melt electrolysis have been
investigated [511–516]. Nevertheless, obtaining a homogeneous cell dispersion within the
vascular beds is still a huge task [516]. None of the methods listed so far enables a selective
and accurate positioning of cells within a 3D-like architecture. In addition, the manufac-
ture of multilayer functional blood vessels with these techniques necessitates a compli-
cated multi-stage production process in which each layer demands a specific maturing
time. Overcoming these main disadvantages has prompted the recent adoption of 3D bi-
oprinting technology, which offers unparalleled benefits [517]. This new technology facil-
itates the production of cellular 3D configurations with various layers in a single manu-
facturing stage, thereby reproducing the inherent hierarchical complexity of vascular tis-
sue [157,518]. Pneumatic extrusion-based bioprinting appears to be the most versatile bi-
oprinting strategy to generate large vessels characterized with centimeter-sized tubular
structures. In this process, the bioink is extruded through a nozzle using a pneumatic sys-
tem and applied layer by layer to a mounting substrate. A wide range of bioinks can be
used in this technology. The most employed kinds of bioinks are hydrogel-based, water-
Cells 2024, 13, 1638 42 of 87
swollen polymer structures whose formulation can be tailored to mimic the ECM envi-
ronment of biological tissue. Nevertheless, the weak mechanical characteristics of hydro-
gels restrict their usability for vascular tissue engineering. The use of a synthetic material
that acts as a scaffolding could be a potential solution [100,115]. For example, three-lay-
ered vascular scaffold has been generated with PCL as a supporting material [519]. Be-
tween two PCL layers, a layer of bioink comprising cells and 3% sodium alginate has been
printed to provide proper structural support [80]. Similarly, poly(ethylene glycol)
tetraacrylate (PEGTA) has been utilized as a carrier system for a cell-responsive bioink
consisting of GelMA and sodium alginate [520]. The bioinks utilized in these studies, nev-
ertheless, are not sufficiently representative of the native ECM constituents. The main
components of the ECM like collagen, elastin, microfibrils, proteoglycans, GAGs and sev-
eral growth factors are indispensable for the preservation of the structural integrity of
large vessels in tissue-engineered vasculature [521,522]. Bioinks on the basis of dECM
have recently been established. The process of decellularization makes it possible to retain
the native microenvironment of the vessels established by the ECM, which encourages
cell growth and a non-immunogenic tissue, while cellular and nuclear components—with
a special emphasis on DNA and RNA—are removed from the tissue [523–526]. This strat-
egy therefore permits the yielding of a bioink consisting of biochemical features existing
in the natural surroundings. The development of a novel bioink made from dECM and
natural hydrogels has been presented, capable of replicating large vascular substitutes
that meet the requirements of the form, functionality and integrity of natural tissue [527].
This strategy brings benefits of the 3D bioprinting and decellularization processes to-
gether. After optimizing the decellularization protocol, the resulting dECM has been in-
tegrated into a bioink whose composition ensures printability and thus overcomes one of
the biggest hurdles of extrusion-based 3D bioprinting [157,528,529]. Finally, the biocom-
patibility of the bioink and cell penetration have been verified by observing cell growth
in 3D-printed constructs over an extended period of time [527].
Moreover, the tumor vasculature is both in structure and function aberrant relative
to the vascular system of healthy tissues. Tumor blood vessels become untight and tortu-
ous, marked by the presence of malformed ECs, detached or displaced pericytes and in-
complete basement membranes. The impermeable vessels, along with the high compres-
sive pressure from condensed cancer cells, compromise the blood perfusion of the tumor
mass and result in interstitial high blood pressure, hypoxia, and acidosis, which have been
demonstrated to ease the infiltration of cancer cells, interfere with drug distribution, and
lead to immune cells invading with cytotoxic capabilities [530]. These circumstances high-
light the relevance of delineating abnormal characteristics of tumor vasculature in in vitro
cancer models. Interaction between tumor and vessels yielded abnormal alterations in
preformed blood vessels due to the surrounding of inflammatory breast cancer [531]. Vas-
cular abnormalities have been characterized by evaluating endothelial shape, cell–cell
connections, matrix porosity, endothelial confluency and permeability [532]. The ap-
proach of quantification enables the investigation of tumor–vessel interfaces and eases the
upcoming trend of vessel normalization, where the organization and functioning of tumor
vessels is adjusted to optimize the effectiveness of cancer therapies. The congenital tumor
vascular arbor, nevertheless, is more intricate than the EC-lined cavity channels engi-
neered in recent reports. Therefore, additional work is required to decipher and evaluate
tumor vasculature aberrations. With recent progress in immunotherapies demonstrating
encouraging anti-cancer effectiveness, there is a pressing demand for vigorous research
frameworks that can fully uncover the interplay between tumor and immune system. Be-
sides the tumor-infiltrating immune compounds that have been demonstrated in various
PDO models [26,44], the involvement of peripheral immune communities residing in the
vasculature deserves attention [533]. With the help of microcirculation meshworks deliv-
ered through bioprinting tumor-on-a-chip technology, peripheral immune cells can be de-
livered through perfusion, allowing the synchronous replication of tumor-infiltrating and
peripheral immune factors in a person-specific TME. This concept supports basic
Cells 2024, 13, 1638 43 of 87
immuno-oncology research, enables efficient therapeutic combination screenings and
opens the door to the development of accurate immuno-oncology.
Tumor metastasis models are fundamental for unravelling the intricate mechanism
of metastasis of cancer to remote organs, which accounts for the majority of human cancer-
related mortalities. Multiple frameworks concentrating on the extravasation mechanism
have been established using replicas of the human microcirculation, cancer cells circulat-
ing intravascularly and immune cells [534]. A high-resolution imaging technique facili-
tates the visualization of the extravasation mechanism to examine the fundamental regu-
latory mechanism [535]. However, the extravasation of cancer cells is a single step in can-
cer metastasis, which is a dynamic and multifaceted phenomenon involving surrounding
the ECM invasion of cancer cells from the primary tumor mass, intravasation, vascular
spread, extravasation and dissemination to distant target organs [536]. To identify the crit-
ical stages of metastasis, it is of great importance to connect the primary tumor and its
potential metastases through a microcirculatory system. The integration of 3D bioprinting
to generate this multi-organ-on-a-chip structure has been shown [537,538], but still re-
quires further research efforts. Thus, the complete elucidation of the complicated multi-
organ crosstalk during cancer metastasis requires further investigations, which may re-
quire the establishment of new techniques [538].
7. Combination of Organoid Bioprinting with Partially Bioprinted Organs-on-a-Chip
Approaches
As organoids lack the flow-driven cultivation of organ-on-a-chip, this weakness can
be overcome by using a combined approach called organoids meet organs-on-a-chip sys-
tems. Moreover, organoids or assembloids and organoid-on-chips will also provide new
platforms for the analysis of collective cell performance in settings of organ-to-organ in-
teractions and host–pathogen interferences [539–541]. Although there are numerous ad-
vantages to utilizing organoids, many present systems lack the cross-tissue cell–cell inter-
action that can induce and sustain collective cell performance in the in vivo milieu. For
instance, the geometric boundary can change the cell stage via mechanochemical feed-
back. In embryonic intestinal tissue, such a boundary is established initially by mesenchy-
mal condensation forming under the intestinal epithelium [542]. It was found that co-cul-
ture of the intestinal organoid with fibroblasts in a 3D collagen I scaffold replicates the
microenvironment of the intestinal stroma and allows the investigation of interactions be-
tween epithelial cells and fibroblasts [543]. In the future, the co-culture of fibroblasts, im-
mune cells and/or neuronal cells within organoids embedded in Matrigel or artificial scaf-
folds, which is currently being established, will expand the study of collective cell behav-
ior in a tissue-authentic environment.
Cells expanded in organs-on-a-chip cultures have been found to upregulate their
functions through maximizing mass transfer and reducing shear stress in the perfusive
soluble microenvironment, bringing them one step nearer to a true natural tissue [544–
546]. A recent experiment demonstrates how the presence of fluid flow facilitates the mat-
uration of renal organoids and their vascularization in vitro [547]. Physical limitations
have been built into the organoid surroundings, and the intestinal cells self-organized into
crypts of the identical size when delineated by artificial scaffolds [381]. Simultaneously,
they surmounted the impenetrability of cystic organoids and the elimination of cell debris
through the generation of a permeable culture of mini-intestines in which the cells are
positioned to generate tubular epithelia and have a similar spatial organization to the tis-
sue in vivo [548]. Future investigations could use the stage to examine the development
of different tissues or the etiology of a variety of diseases, thus enabling the identification
and preclinical testing of drugs. In all these techniques, organoids or assembloids are pre-
ferred for bioprinting. There are fewer arguments for the use of organs-on-a-chip systems
due to the weaknesses in mimicking the native in vivo environment of organs. In addition,
organ-on-a-chip systems can only be partially printed [549]. However, 3D bioprinting can
be utilized to produce not only microfluidic chips made of materials like resins and
Cells 2024, 13, 1638 44 of 87
polydimethylsiloxane, but also biomimetic tissues derived from bioinks like cell-laden hy-
drogels. The combination of both techniques, such as bioprinted organoids and bioprinted
organ-on-a-chip systems, holds great potential for improved organoids or assembloids-
based model systems in a dynamic environment, such as a fluid-flow condition.
8. Four-Dimensional Printed Materials for Cells, Hydrogels and Organoids
Four-dimensional printing is a process for producing a model system from one or
more materials that can be transformed from a 1D strand into a pre-programmed 3D
shape or from a 2D surface into a pre-programmed 3D shape and is also capable of trans-
forming between different dimensions [550]. These transformations are accomplished, for
instance, by heating, light or swelling in a liquid, electrochemically and by programming
different degrees of sensitivity, e.g., for swelling, in various areas of the designed geome-
try [551]. These techniques offer flexibility and dynamic behavior for structures and sys-
tems of all sizes and open up new possibilities for incorporating programmability and
simple design principles into non-electronic materials [552]. In biological systems, the de-
velopment of the macrostructure of the engineered material is frequently altered by the
soaking of the hydrogels, as the engineered material is grown in aqueous solutions to
maintain the cells in a hydrated state. The swelling characteristics of a bioprint are deter-
mined by the intrinsic polymer solubility, the crosslinking degree and the degree of het-
erogeneity of the structure [553,554]. For instance, the frequently employed hydrogel PEG
expands substantially due to its highly hydrophilic, water-soluble characteristics, whereas
the amphiphilic polymer Pluronic (a PEG-polypropylene glycol (PPG)-PEG triblock pol-
ymer) can only absorb water to a limited extent because of its hydrophobic nature [555].
The degree of swelling of hydrogels also reduces as the level of crosslinking rises; whereas
the crosslinking level can be controlled to regulate swelling, the rheological characteristics
of the hydrogel are influenced at the same time [556]. For heterogeneous bioprints with
several different material types or intricate geometries, limited spatial or non-isotropic
swelling can lead to significant geometric variations after printing. These temporal fluc-
tuations in the biologically printed structure can be used to create what are known as “4D”
bioprinted materials [557]. For instance, a simple 3D-printed structure can transform into
a more complex structure over time [558]. The principle of printed active composites
(PAC) has been established, which can transform a printed film into a complex structure
using the shape memory effect [559]. Global research into 4D printing has grown expo-
nentially. There are many definitions of the 4D printing process. An early definition
claimed that 4D printing is just 3D printing over time [558,560,561]. The definition that
clearly specifies 4D printing, nevertheless, states that 4D printing is the evolution of the
shape, properties and performance of a 3D-printed structure over time as it is exposed to
heat [562,563], light [564,565], magnetic field [566], pH [567], water [561,568] and other
similar things. Another definition says that 4D printing is the creation of scaffolds that
alter their form as they leave the 3D printer. These objects self-assemble when they en-
counter water, heat, air, etc., due to the chemical reaction of the material used. Four-di-
mensional printing is a combination of a 3D printer, a smart material and a correctly pro-
grammed arrangement [569,570]. In 4D printing, various metamaterial structures are
formed when the environment undergoes a change. Currently, most of the research in 4D
printing technology focuses on the ability of 4D-printed materials to alter their form, such
as by stretching, bending, curling and twisting. In biological and medical fields, the 4D-
printed structures can alter their physical/chemical properties like stiffness or density. In
addition, they display various phenomena, including shape memory effects and shape
transformation [571], which, in turn, may impact cellular functions. The shape memory
effect is a mechanism by which a system/structure can remember a particular shape and
shift from one shape to another; for example, from the original shape to a programmed
shape, in a planned way facing external cues. Shape shifting is a natural process in which
a system/structure can change its appearance from one shape to another due to external
influences. Compared to 3D printing, 4D printing offers several advantages, such as rapid
Cells 2024, 13, 1638 45 of 87
growth of smart and multi-material materials, more flexible and malleable patterns and
more applications for 4D or 3D printing. In addition, 4D printing offers higher efficiency,
quality and performance compared to conventional techniques, as the 4D-printed struc-
tures can self-improve their characteristics and performance.
8.1. Key Drivers of 4D Printing
Four-dimensional printing is based on primarily five parameters. All these five pa-
rameters need to be accounted during the process of 4D printing. These five parameters
comprise the AM process, the printing material, the stimuli, the mode of interaction and
the type of modeling [570]. The first issue is the AM technique employed for printing. The
AM procedure permits the manufacture of print media from the digital data received by
the computer without the need for an intermediate device [550]. Several AM techniques
exist like SLA, 3D printing with nozzles (3DP), selective laser melting (SLM), selective
laser sintering (SLS), fused deposition modeling (FDM), direct ink writing (DIW), electron
beam melting (EBM) and similar, and virtually all of these can print 4D media, provided
the material being printed is suitable for the type of printer [572]. The next element is the
material utilized for the print, which needs to react to the impulses as it is deposited layer
by layer. These materials are also referred to as programmable or smart materials [282].
The nature of these smart materials dictates the type of stimuli to be applied, and the re-
action of these materials to the various stimuli defines the material’s capacity for self-
transformation. The third factor is the type of stimuli involved in 4D printing. The stimuli
employed may be physical, chemical or biological in nature [571]. Physical stimuli com-
prise light, humidity, magnetic and electrical energy, temperature and UV light. Chemical
stimuli encompass chemicals, the pH value, the utilization of oxidizing substances and
reducing media. Biological stimuli involve enzymes and glucose [573]. When the stimulus
is implemented, physical or chemical modifications like relaxation of the stress, move-
ment of the molecules and phase modifications in the network are induced, which cause
the structure to distort. The fourth and fifth features are the mode of interaction and its
mathematical modeling [550]. When an intelligent material is subjected to a stimulus, not
all materials can experience the intended alteration. An interplay mechanism like mechan-
ical stress or physical force must be supplied to design the sequence of form modification.
Following the provision of the interaction mechanism, mathematical modeling is neces-
sary to schedule the time, during which the stimulus affects the smart matter [550].
8.2. General 4D Printing Laws
Three laws of 4D printing have been defined that determine the form-changing prop-
erties of all 4D-printed textures [574]. These laws allow a better comprehension of the
physics underlying the shape-alteration capability of 4D-printed patterns. They are for-
mulated as follows: The first law says that all form-alteration phenomena like the wind-
ing, rolling, twisting, bending, etc., of 4D composite patterns are caused by the mutual
extension between active and passive components [575]. The second law says that four
physical drivers are responsible for the capacity of all multi-material 4D patterns to trans-
form their form: diffusion of mass, thermal dilation, molecular transformation and organic
outgrowth [576]. All these features result in a mutual stretching between active and pas-
sive substances, which causes a modification of the form when a stimulus is applied. The
absorption or adsorption of irritants, such as water or ions, leads to a modification of the
mass of the network. This transportation of material ultimately results in a relative exten-
sion of the material and therefore a distortion of the form. The alteration in mass can also
be caused through thermal, electrical, chemical, or light cues [550]. Thermal expansion can
lead to distortions in the physical structure, as the average spacing between atoms and
molecules grows or shrinks with rising or falling temperatures, leading to a relative dila-
tion [577]. Thermal extension can also arise when electrical, light and UV signals are ap-
plied, as these may modify the temperature of the texture. In structures where the mass
and temperature are unchanged, there may be relative expansion through molecular
Cells 2024, 13, 1638 46 of 87
conversion. In these instances, electrical fields, magnetic fields, light or mechanical forces
can affect a molecular conversion [576]. For instance, when an electric or magnetic field is
imposed, the dipoles in the substance orient themselves in the direction of the imposed
field, leading to a conversion of the molecules. Likewise, applying mechanical force to
polymers forces the polymer chains to orient in a particular orientation, and irradiating a
photosensitive material with UV light causes it to convert from trans to cis [578]. Organic
growth refers to the gain in length and weight of a living organism within a specific time
frame. The increase in weight and elongation causes a mutual extension of active and pas-
sive substances, which leads to form-modifying characteristics. In living organisms, or-
ganic growth can usually be induced through electrical impulses [579]. However, in ad-
dition heat, water, pH level and mechanical stress can also be utilized. Organic growth is
employed to characterize the deformation response of cells, scaffolds, tissues, organs and
stents, designed using 4D bioprinting [579] Organic growth is employed to characterize
the deformation response of cells, scaffolds, tissues, organs and stents, designed using 4D
bioprinting [550]. The third law of 4D printing says that “the temporal deformation char-
acteristics of nearly all multi-material 4D-printed patterns are governed by two kinds of
temporal constants”. These constants may be the same or vanish based on the type of
input stimulus and what material is utilized for 4D printing [580].
8.3. Material Types Used in 4D Printing and Generation of Tissue-like Constructs
and Organoids
In 3D printing, latest advances have allowed materials to be positioned more accu-
rately and with more flexibility, which has greatly benefited 4D printing [581]. The mate-
rials employed for 4D printing are usually known as smart materials, because they can
alter their characteristics as time passes (Figure 7) [582]. These materials can react to out-
side impulses and have characteristics including self-organization, self-repair, form reten-
tion and self-sustainability [213]. In addition, 4D printing involves not simply materials
that can transform their form but also undergo color alterations upon exposure to UV or
visible light [574].
Figure 7. Material types employed for 4D bioprinting.
8.3.1. Materials React to Moisture: Hydrogels
Materials that react to moisture or water have gained a lot of interest because of their
wide array of applications. These materials are also referred to as hydrogels, as they have
4D printing
materials
Hydrogels
(Materials reactive to
moisture)
Electro-responsive
(Materials reactive to
electric energy)
Piezoelectric
(Materials reactive
to mechanical
stress)
Thermo-responsive
(Materials reactive to
heat or temperature)
Photo-responsive
(Materials reactive to
light)
Magneto-responsive
(Materials reactive to
magnetic energy)
pH-responsive
(Materials reactive
to pH alteration)
Cells 2024, 13, 1638 47 of 87
an exceptional capacity to respond to water or humidity. These are a class of 3D polymer
chain meshes that are formed by networking and can extend by up to 200% of their initial
volume after encountering humidity. Hydrogels also have a high compressive strength,
as different textures have been designed with hydrogels that can be wrinkled, flexed,
stretched and geometrically expanded. These are extremely biocompatible and simple to
print on when they are written with direct ink [583]. The only issue is their slowly revers-
ing response, which means that drying and shrinking takes several hours. Overcoming
this problem depends on programming the hydrogels so that their swelling is enhanced
by anisotropy. Cellulose fibrils have been paired with the hydrogen ink and these have
been oriented through the evolution of shear forces generated by the physical contact of
the print plate and the hydrogel ink [568]. This orientation made the transverse swelling
four times larger than the longitudinal swelling, which enabled the programming of the
printed 4D texture. Another possibility is to limit the hydrogels in a single direction with
rigid materials, causing an anisotropic expansion of the hydrogel [581]. Films of stearoyl
ester (CSE) cellulose have been prepared, and these hydrophobic films exhibited a more
accurate and rapid reaction than the previous films [584]. Usually, hydrogels are added
to water and take up the water until their saturation level is achieved. The problem with
this mechanism, however, is that it restricts the capacity of hydrogels for intermediate
regulation. This problem can be solved through regulating the temperature of the aqueous
medium. This was seen when producing the microgrip compound utilizing (poly-N-iso-
propylacrylamide-co-acrylic acid) pNIPAM-AAc hydrogels [585]. Reverse actuation is
possible when the temperature of the water in which the hook has been dipped is altered.
The 4D material has also been printed with alginate/pNIPAM ICE gel inks [586]. The use
of hinge structures to avoid undue swelling has been proven. A self-pleating pattern has
been manufactured from PolyJet printers that ceases to pleat at specific angles that were
pre-programmed to prevent over-swelling [558].
8.3.2. Materials React to Temperature: Thermo-Responsive
These objects are smart materials that react to heat or temperature signals. The vari-
ations in the form of these materials in response to thermal impulses are primarily at-
tributable to two mechanisms, such as the shape change effect (SCE) [587] or the shape
memory effect (SME) [588]. SME refers to the transformation of a shaped (plastic) material
into its initial form through external stimulation [589]. Smart materials displaying the
SME effect are referred to as shape memory materials (SMM) and are categorized as shape
memory alloys (SMA), shape memory ceramics (SMC), shape memory hybrids (SMH),
shape memory gels (SMG) and shape memory polymers (SMP) [590]. SMMs are classified
into one-way, two-way and three-way (or multiway) materials based on the degree of
shape transitions (Figure 8).
Cells 2024, 13, 1638 48 of 87
Figure 8. Shape memory materials (SMMs). A one-way SMMs cannot regain its initial shape after
being deformed. In contrast, two-way, three-way and multiple-way SMMs can restore their initial
shape. Two-way SMMs can switch between the two shapes. Three and multiple-way SMMs can
return from a temporary shape to the initial shape via one intermediary shape or multiple interme-
diary shapes following deformation.
With one-way SMMs, the original form cannot be restored following deformation,
while with two-way and three-way SMMs, the initial form can be returned to a temporary
form after deformation through an intermediate shape [591]. SMMs can exhibit SCE
alongside with SME, according to the ambient requirements. Among the different types
of SMM, the SMPs are frequently employed as they can be printed with ease. The SMPs
can regain their initial form after being deformed with the appropriate irritant [592]. The
SMPs feature a typical glass transition temperature (Tg), generally exceeding the temper-
ature where they are normally worked. Beyond Tg and under certain thermal and me-
chanical constraints, SMPs undergo programming and, as they cool, take on a transient
shape that is devoid of any outside stress. When the temperature is increased beyond Tg
once again, they take on their initial shape [593]. Below the Tg, the internal energy of the
polymer chains is minimal, and they are unable to move freely, resulting in the material
becoming glassy and stiff [594]. Above the Tg, however, energy is supplied to the polymer
chains allowing them to move, resulting in the material appearing like rubber and prone
to distortion and tampering [595,596]. An SMP sphere has been produced with the SLA
printing process, and the sphere could switch between a flat plane and its initial form with
high sustainability [597]. It has been demonstrated that an SMP may be pre-programmed
in FDM printers through heat [593,598]. The SMPs have been adapted to utilize their dis-
tinctive characteristics for printing applications, that is, thermoset and thermoadapted
SMPs [599–603]. The SMPs exhibit two or three interim transitions, and it is even possible
to keep an intermediate state, which is equally stable. Another SMM commonly utilized
in 4D printing is SMA, which is capable of altering its shape depending on the tempera-
ture fluctuation [604]. SMAs exhibit a typical temperature, referred to as the transfor-
mation end temperature, beyond which they display a high yield strength and a high
Young’s modulus, which means that they are superelastic when above this temperature
[604]. Nitinol (nickel-titanium) has been found to be the widely utilized SMA because of
its favorable SME characteristics, its high ductility and toughness, and its robust cyclic
characteristics, which renders it to be more biocompatible and amenable to actuation [605–
607]. A Ni-Mn-Ga-based SMA has been examined for printing 4D components with the
binder jetting technology, where the printed components displayed a reversed form alter-
ation upon cooling and heating [608]. Cu-based SMAs have also been investigated for 4D
One-way SMMs Two-way SMMs Three-way or multiple-way SMMs
Stimulus triggered
Stimulus triggered
Stimulus removed
Stimulus triggered
Stimulus removed
Stimulus triggered
Stimulus removed
Intermediary shape 1
Initial shape Initial shape Initial shape
Temporary shapeTemporary shapeFinal shape
Stimulus
triggered
Stimulus
removed
Intermediary
shape 2
Cells 2024, 13, 1638 49 of 87
printing because they can efficiently endure post-printing operations and are inexpensive.
Although they are not as favored as SMAs because of their limited ductility, several pieces
of research have been carried out involving these Cu-based SMAs [609–611]. Another type
of SMA on the basis of iron (Fe) is also being explored for 4D printing utilizing the SLM
printing method. They are inexpensive and feature pseudoelastic elongations [612]. SMAs
have been mainly studied for their use in the biomedical sector, including surgery, ortho-
dontics and physiotherapy [613]. The key restriction to the use of SMAs in 4D printing lies
in their high expense. In addition, SMPs are lighter, more elastic, biocompatible and re-
quire lower energy consumption compared to SMAs [614,615]. To address the constraints
of SMP and SMA, shape memory composites (SMC) have been designed through the in-
tegration of SMP with SMA or SMP along with a strengthening fiber [616]. The strength-
ening fiber may be a long or short fiber, nanoparticle or nanofiber with elevated mechan-
ical characteristics and high deformability [616]. In addition, SMHs represent the intelli-
gent materials created by the combination of SMAs, SMPs and hydrogels. These can react
to temperature, pressure and a number of different inputs simultaneously. In SCE mate-
rials, the deformation is linear to the exerted stimulus or the deformation can be charac-
terized as changing from one extreme state to another [587]. When the material reacts
thermally, SCE occurs in the two-layer constructions. The structure flexes when the load
is applied, but the interface between the layers stays the same. A graphene-based network
has been produced and demonstrated to transform into a flat plate when heated and re-
gain its initial cylindrical form upon cooling [617]. Such drastic alterations in form neces-
sitate a strong variation in temperature.
8.3.3. Materials React to Light: Photo-Responsive
Light also works as an implicit impulse for the distortion of smart materials. When a
region of a smart material that interacts with light, which is referred to as a photosensitive
material, is subjected to light, it absorbs the light, causing the material to heat up. Heat
acts as a form of impulse for the shaping of intelligent materials, resulting in a form alter-
ation of the photosensitive material. A sequential self-folding structure has been demon-
strated in which the light is absorbed by the joints, and these are subsequently heated,
causing a transformation in form [618]. The speed at which the heat is taken up by the
joints varies according to the light source applied and the joint color. The light can be
employed in a different manner to evoke a deformation within the photo-responsive sub-
stance. A light-sensitive chromophore can be incorporated in certain areas of a polymer
block (gel) so that just these areas are distorted when light falls on the pattern [564]. Ad-
ditional work demonstrates that UV light (weak) and visible light have been applied to
distort the 4D architecture [619].
8.3.4. Materials React to Electric Energy: Electro-Responsive
Electricity acts as an indirect impulse in the same manner as light, as it has been
demonstrated to induce heating because of the resistance of the substance it passes
through. For this reason, materials that deform because of their reaction to electric current
are referred to as electro-sensitive materials. An engineered muscle has been manufac-
tured from a blend of ethanol and silicone elastomer. Current conducted across the muscle
leads to the vaporization of ethanol, thereby expanding its volume, which inevitably de-
forms or stretches the muscle [620]. The absorption or desorption of water in polypyrrole
(PPy) can be guided through electricity, and this concept has been utilized to fabricate
microrobots (origami) of PPy. The absorption of moisture when the robot has been held
in a humid surrounding caused a tension that propels the head of the robot to the front.
The tail of the robot trails the head when desorption takes place and there is no tension
[621].
Cells 2024, 13, 1638 50 of 87
8.3.5. Materials React to Magnetic Energy: Magneto-Responsive
The magnetic field or magnetic energy acts as an indirect impulse that can induce
distortion in smart materials. The materials employed for printing 4D structures based on
their distortion reaction to magnetic energy are termed magneto-responsive materials. By
using magnetic nanoparticles in microgrippers that were printed from hydrogel, a state
of remote controllability has been introduced. When a magnetic field has been placed on
the printed pattern, it started to display a response that can be managed remotely [585].
The 4D-printed structures utilizing magneto-responsive media have enormous potential
in the area of metal and polymer printing, with the sole disadvantage that the print size
needs to have a low mass so that it can be influenced through the magnetic field. Magneto-
active 4D printing is based on magnetic field induced systems. The printing materials,
such as resin, powder or filament need to contain magnetic field sensitive features/com-
ponents, which are generally fillers, that are activated through an outside magnetic field
to exhibit the 4D phenomenon. The first key stage in the creation of 4D structures with
3DP is therefore the process of adapting the printing materials through the inclusion of
active elements. The most frequently deployed magnetoactive filling materials are car-
bonyl iron powders (CIPs), iron(II, III) oxides and Fe-Nd-B micro/nanoparticles [622–626].
Not all these magnetic fillers can still be applied in all 4D printing processes which is
mainly because of the dimensions of the fillers. Nevertheless, there are alternative printing
processes, like DIW, in which nanometer-sized to micron-sized filling materials have been
employed with great efficiency [627–631]. The fillers have been designed to improve the
ability to re-extrude compound filaments for FDM [632,633], the manufacture of compo-
site or surface-decorated (with nanofillers) micropowders for SLS [634] and high strength
in fluid resin for SLA/DLP [635–637]. Pure SLA is amended to its derivatives, such as di-
rect laser processing (DLP), micro-continuous liquid interface production (μCLIP) or two-
photon polymerization (2PP). However, all variations are still reliant on the light-driven
transition of liquid resin [135,638–640].
8.3.6. Piezoelectric Materials Distort upon Force
Piezoelectric materials are employed in 4D printing purposes as they are able to dis-
tort under the impact of a mechanical force [641]. As a type of smart material, piezoelectric
materials can create an electrical current when they are exposed to mechanical stress that
inevitably cause alterations in the structure, because the charge can induce deformation.
These piezoelectric materials can therefore alter their geometry when an electric current
is supplied to their surface. These materials consist of crystals or ceramics that exhibit a
certain crystal structure, like quartz or barium titanate. The effect of generating electrical
charges due to the action of a mechanical force is termed piezoelectricity. The piezoelectric
effect is due to the orientation of the crystal structure, which creates an electrical charge
while the material is distorted [642].
8.3.7. Materials React to pH
These are intelligent materials that react to the pH value and can alter their form and
volume depending on the pH level [643,644]. The form modification in reaction to various
pH values renders them appropriate for 4D printing applications. Polymers that react to
pH have been utilized for 4D printing, like polyelectrolytes, as they can absorb or release
protons when the pH is altered due to an ionizable side chain. Upon the liberation of a
proton, the polymer string expands due to electrostatic repulsion, leading to a defor-
mation of the network, and when a proton is absorbed, the network neutralizes. Polyelec-
trolytes include polycations or polybases like ammonium salt as a functional chain and
polyanions or polyacids, such as carboxyl or sulfone moieties as ionizable side groups.
The side chains donate the proton at higher pH values (stretching) and absorb the proton
at lower pH values (neutralizing). The functional group, however, donates the proton at
lower pH values and absorbs the proton at higher pH conditions [645–649]. pH-
Cells 2024, 13, 1638 51 of 87
responsive materials can be used in drug administration [647,650], soft robots, actuators
[643], valving, biocatalysts and in stabilizing of colloids [567,648]. Smart materials can re-
act to a single or two kinds of inputs. For example, SMPs can be made to interact with
temperature, light and electrical energy, while compound composites can react to various
irritants.
8.3.8. Maturation of Hydrogels via Physical Stimuli Toward Tissue-like Constructs
and Organoids
The maturation of hydrogels and cells into tissue-like assemblies has also been
speeded up by the introduction of physiologically important external factors. In the tissue
engineering of cartilage tissue, the mechanical pacing of structures under pressure en-
hances ECM production and encourages cartilage cell diversification [651]. In vitro mod-
els of cardiac tissue take advantage of electrical pacing that encourages differentiation
along a cardiomyocyte lineage and helps synchronize the heartbeat [652,653]. It is antici-
pated that continued evolution of bioink materials that facilitate these types of post-print
extrinsic inputs will result in more “mature” engineered tissues. For instance, the utiliza-
tion of injectable conductive hydrogels as bioinks in combination with the use of external
electrical stimuli following printing may provide a bespoke approach for cells that need
electrical signaling for proper performance. By taking into account the physiological fac-
tors specific to the target tissue type, bioprinted designs can ripen to get more similar to
natural tissue.
9. Cell Alignment in Printed Scaffolds
The alignment of cells is critical for cellular functions, such as the generation of forces
and cellular motility, like cancer cell invasion, or developmental processes, such as organ
formation. Moreover, the functionality of organs relies on alignment-driven functions. For
instance, the electrical and mechanical characteristics of the heart [654] and the multinu-
cleation of muscle fibers during the creation of myotubes [655] need, for musculoskeletal
tissue, a high amount of cellular alignment to carry out key cellular roles. In addition,
aligned cells inside a highly organized, anisotropic ECM set off a series of events that are
crucial for defining the function of the tissue [656,657]. The alignment of cells appears to
contribute to several cellular behaviors, including the reorganization of the cytoskeleton,
nuclear gene expression, and rearrangement of the ECM scaffold. The alignment and elon-
gation of cells in the direction of anisotropic and aligned topographies are important phe-
nomena of cellular contact guidance and are seen in multiple cell types. Hence, a question
arises whether there exists a universal mechanism behind cell alignment [658]. The most
commonly acknowledged model of cell alignment is topographically-driven orientation,
which proposes that anisotropic topographies constrain the growth of focal adhesions and
actin stress fibers laterally, thereby promoting anisotropic force generation, cellular
stretching and alignment. There are certain circumstances in which alternative or comple-
mentary mechanisms of cell alignment seem to come into effect. These examples involve
the cases of certain cell types, like amoeboid cells and neurons, and particular topogra-
phies. Moreover, the actin cytoskeleton is involved in regulating topographically based
cell alignment, highlighting the importance of elucidating the contribution of other cyto-
skeletal components. The understanding of cell alignment is critical for identifying the
function of cellular contact guidance in healthy and diseased conditions.
Support-assisted bioprinting is employed to bioprint a hydrogel with embedded
C2C12 cells into a lattice structure [659]. The purpose of employing a lattice structure is,
firstly, to assess the structural accuracy of guided bioprinting. Secondly, the lattice pattern
is utilized to imitate the angular variations of fibers in various layers. In this approach, a
secondary material is incorporated as a mold for the encapsulation of a primary material.
The secondary material can be fully taken out. The bioink consists of 10%w/v GelMA and
2%w/v alginate [659]. The purpose of support bioprinting is to ensure the structural rigor
of the primary material before the ultimate crosslinking of the bioprinted structure. It can
Cells 2024, 13, 1638 52 of 87
be assumed that the cell alignment parallel to the printed hydrogel rods of the lattice struc-
ture can be realized. Extrusion-based bioprinting can be utilized to control cell orientation
with the help of a predefined extrusion path. This allows the angle of cell alignment to be
specified on various layers, in a similar way to native myocardial features. This novel bi-
oprinting strategy has been shown to achieve macroscale alignment of cells, resulting in
an 80% alignment of cells falling within a 15° orientation. Moreover, the printing strategy
demonstrated regulated cell alignment corresponding to the different angle modifications
on different planar levels.
The orientation of the cells inside the printed substrate can be accomplished through
magnetic tagging of MSCs and HUVECs. The magnetization of cells using standard fluo-
rescent MNPs from Chemicell (100 nm) has been carried out following a conventional
approach. HUVECs have been magnetized with nano-screen MAG/R-PAA nanoparticles
bearing a red fluorescent tag, whereas MSCs have been magnetized with nano-screen
MAG/G-PAA nanoparticles bearing a green fluorescent tag. The use of a magnetic frame-
work can orient them to imitate the vascularization of osseous scaffolds [660]. A new,
multi-purpose and user-friendly approach has been developed to promote controlled 3D
sowing of cells by magnetic guidance. The simple pulling of cells charged with magnetic
nanoparticles across an external magnetic field has previously been published and re-
sulted in solutions that are beyond the scope of any other technology [660–665]. Specially
engineered magnetic frameworks are employed that are featured by strong magnetic gra-
dients on a short scale (100–200 μm) that can align and capture the magnetized cells on
the selected face of the framework fibers. Such local magnetic structuring constitutes a
practical way to build 3D cell structures with a controlled structure at the microscale. As
principal proof of this exceptional capability, a well-defined separation of two cell popu-
lations, specifically MSCs and HUVECs, has been achieved on the confronting sides of the
magnetic osteogenic framework fibers. This cell composition is anticipated to support the
bone microarchitecture restoration with suitable characteristics, especially with regard to
the vascularization of the artificial bone [666,667]. The cells have been magnetically tagged
with MNPs, whereas the frameworks have been modeled and manufactured on the basis
of sophisticated magnetic materials by blending bioresorbable Fe-doped hydroxyapatite
(FeHA) with PCL [668]. In concrete terms, the 3D frameworks have been created through
injection/extrusion and laying down the fibers in specific orientations in line with the spec-
ified laying template [669,670]. The nanocomposite pellets have been exposed to a tem-
perature of in the range of 110 to 130 °C in a cartridge unit attached to the flexible arm of
a 3D bioprinter. The magnetic force exerted on the cells is a function of their magnetization
and the regional gradient of the magnetic field [671]. To create analogous conditions for
cell handling and adhesion to the scaffold, it is essential to obtain analogous magnetically
actuated forces for most cells. Performing standard magnetic mapping of magnetized cells
is a difficult challenge, as it is hard to keep cell-friendly environments in magnetometers
or susceptometers. Therefore, an assay focusing on cell locomotion in an imposed mag-
netic field has been established. Cell motility in reaction to magnetic guidance has initially
been assessed in a qualitative manner by utilizing a cylindrical NdFeB permanent magnet
exhibiting 1.2 T magnetic remanence. The magnet has been positioned under the base of
the cultivation plate and almost all cells have been magnetically captured, with a marginal
quantity of cells staying outside the magnetic field.
Another method of distributing the bioprinted cells at the micro-architecture level is
an approach using the acoustophoresis principle. The physics of ultrasound-assisted bi-
oprinting (UAB), exploiting the principle of acoustophoresis can be exploited to orient
MG63 cells inside single- and multilayer after bioprinting of alginate constructs using ex-
trusion [663]. The cells have been oriented orthogonally and parallel to the printed fila-
ments, thereby imitating cellular anisotropy in tissues like ligaments, tendons and cardiac
muscle. In a similar manner, an acoustic excitation mechanism has been utilized to direct
skeletal myoblast cells (C2C12) and HUVECs embedded in GelMA bioink [664].
Cells 2024, 13, 1638 53 of 87
Aligned cells exhibit directionally sensitive mechanical characteristics that affect bi-
ological and mechanical functionality in native tissues. Conventional alignment tech-
niques like casting and uniaxial stretching are unable to completely mimic the intricate
fiber alignment of native tissues in organs such as the heart. Bioprinting is utilized to
guide cell orientation. A 0°–90° lattice pattern was 3D-printed to determine the strength
of the supported bioprinting technique [659]. Changing the angles of the lattice pattern is
intended to replicate the variations in fibril alignment in native tissue, where the angles
of cell alignment change through the various layers. When a cell-hydrogel blend was bi-
oprinted, C2C12 cells showed an alignment in the direction of the printed beams. Cell
alignment is accomplished by inducing structurally stable structures, such as various 0°–
90° structures, and by permitting cells to dynamically reshape the bioprinted structure. A
heterogeneous scaffold has been developed [672], that can be adjusted in terms of gradient
strength, adjustable fiber diameter and pore size. Heterogeneous scaffolds containing ul-
trafine fibers, like 3–22 µm in diameter, can be printed using high-resolution melt elec-
trowriting (MEW) with single-nozzle printing by adapting the printing conditions. The
diameter of the printed fiber is similar to that of cells (tenths of a micrometer), less than
that of traditional 3D-printed scaffolds (100 µm or more) and cell adhesion is very suscep-
tible to fluctuations in fiber diameter. By precisely positioning thick and thin fibers, cells
can be made to stretch in a grid in the expected direction. For example, cell alignment in
a singular scaffold with four regions displayed four different alignments, all separately.
The scaffolds may also be non-uniform in pore size, whereby the proliferation velocity of
the cells in small pores is three times elevated compared to that of the cells within large
pores. In addition, different pore sizes and fibers can be incorporated into a scaffold, ena-
bling cell growth to be guided to various stages by tailoring the scaffold architectures.
This approach usually constitutes a method for structure-induced cell growth to better
imitate the in vivo microenvironment.
Conventional electrospinning has been widely used to produce tissue-engineering
scaffolds [673,674]. Due to its ultrafine continuous fibers, high surface-to-volume ratio and
high porosity, the electrospun scaffold has morphological resemblances to the natural
ECM [675,676]. Nevertheless, the fibers are aligned randomly, which impedes efforts to
regulate the scaffold structure created [676]. While some specialty collectors, such as U-
collectors, parallel plate collectors and rotating tumble drums, are designed to maintain
aligned fibers, it is challenging to manufacture free-flowing structures [677–680]. In an-
other application, 3D printing has been used to create scalable scaffolds for tissue engi-
neering, where the size is typically between hundreds of micrometers and millimeters.
The advantages of traditional electrospinning and 3D printing have been coupled to create
electrohydrodynamic (EHD) printing [672], which facilitates the manufacture of porous
scaffolds with precisely arranged ultrafine fibers [681,682]. The technique has also been
applied to the production of pliable electrodes [683–685]. In comparison to conventional
electrospinning, electrical instabilities, also referred to as the “whip effect”, are avoided
by decreasing the distance between the nozzle and collector during EHD printing [686].
By applying a relative motion between the nozzle and the collector, EHD direct writing
enables high-resolution patterning of the micro/nanofiber [687,688]. Melt-driven EHD is
a physical and environmentally safe printing technique that does not involve organic sol-
vents and avoids the restrictions imposed by toxic residues and their build-up [689]. The
technique is also better at producing scaffolds for clinical tissue healing. There are even
other techniques, such as dielectrophoretic-based bioprinting and rotary jet spinning. For
example, the fabricating of a dielectrophoretic microfluidic device has been described that
employed 3D-printed molds and silver conductive paint [690]. The latter, rotary jet spin-
ning technique has been employed to produce ECM scaffolds and seems to be suitable for
combination with bioprinting. This approach has been seen to decrease the amount of
bacterial contamination of printed scaffolds [691].
Finally, bioprinting is emerging as an important manufacturing tool for tissue engi-
neering by controlling cell orientation through the design of the printing path of material
Cells 2024, 13, 1638 54 of 87
laydown. Spatial orientation of 3D-printed scaffolds, such as blended gelatin-sodium al-
ginate 3D-printed scaffolds, regulates the gene expression profile of pre-osteoblasts [692].
Additionally, 3D-printed graphene-PLA scaffolds enhance the cell orientation of iPSC,
neuronal cells, immortalized fibroblasts and myoblasts and enhance their differentiation
[693]. In addition to orienting cells based on topological cues, graphene is instrumental in
driving cell differentiation, as evidenced by iPSC compulsion to neuroectoderm and the
merging of myoblasts into multinucleated myotubes accomplished through the 100 µm
graphene scaffolds [693]. This research demonstrates the creation of a robust and econom-
ical 3D-printed scaffold with the capability of being employed in multiple tissue engineer-
ing applications and reveals how the scaffold’s microtopography and the characteristics
of graphene synergistically guide cell differentiation.
10. Conclusions, Open Questions and Future Directions
The synergistic use of bioprinting, scaffolds, organoids, organ-on-a-chip and ad-
vanced biomaterials opens a completely new branch of cancer research with the aim to
create tumor models that more faithfully reproduce the TME. However, the use of specific
organotypic tumor models for the purpose of mechanobiological analysis and ultimately
the study of disease mechanisms from a biophysical perspective still faces major hurdles.
Thus far, the mechanical analyses of cancer cells and their environment such as the ECM
scaffold and adjacent cells have been performed in cell culture systems mainly derived
from 2D cell cultures, or even in a capillary system without significant adhesion of the
cells, which surely does not correspond to the natural environment of these cells in tissues.
There are still some key questions that remain open, such as the following. What efforts
can be made to increase the efficiency of growing personalized tumor organoids and gen-
erate patient-specific tumor models for each cancer subtype in the entire patient popula-
tion, incorporating the structural and mechanical features within tumors or in tumor
niches during malignant progression? What are the potential advantages for mechanical
studies of single cells within 3D organoids, which are embedded in natural or synthetic
or hybrid biomaterials, when using 4D bioprinting? Could these models be used to inves-
tigate how these cells respond to stimuli for post-printing modification of in vitro tumor
models? Would it be possible to analyze the temporal course of stimulation? How can 3D
bioprinted or 4D bioprinted tumor models help to improve the prognosis and diagnosis
of tumors and their malignant progression, taking into account mechanobiological as-
pects? How can a combination of clinical imaging and genetic analysis be used to identify
predictive biomarkers and/or mechanomarkers for tumor progression and determine spe-
cific biochemical and physical mechanisms involved in the therapeutic response to heter-
ogeneous medications? Overall, the analysis of mechanical aspects in 3D-printed organ-
oids with vascularization is crucial for the accurate analysis of mechanical drivers or me-
chanical markers for the malignant progression of cancer. There is great potential to iden-
tify general mechanomarkers of cancer progression that are not limited to a specific cancer
type. In addition, there is a unique opportunity to develop mechanically heterogeneous
3D organoids of cancer cells, which may also contain immune cells, such as macrophages,
or lymphocytes, stromal cells and a specific architecture of the vasculature.
10.1. Patient-Specific Modeling of Tumors
Current cancer models largely rely on tumor spheroids derived from immortalized
cell lines because of their ease of proliferation and capacity to withstand intricate fabrica-
tion procedures. Nevertheless, cell lines are prone to accumulate genomic aberrations and
cannot replicate the heterogeneity of tumors in individuals, leading to divergent pharma-
cological reactions. Patient-derived cell collections are preferable for use in precision med-
icine and personalized drug testing. They nonetheless encounter the challenges of limited
specimen resources, poor cell viability during transportation and low organoid retrieval
efficiency. For example, the establishment level is under 30% for certain subtypes of can-
cers [694,695]. Hence, there is an immediate demand for improved primary cancer cell
Cells 2024, 13, 1638 55 of 87
collection methodologies, cryopreservation techniques and enhanced culturing tech-
niques to guarantee the effectiveness of organoid collection in the entire patient popula-
tion and across the full range of cancer subtypes [696,697]. The pairing of tumor organoids
and stromal cells from the identical patient could aid the identification of more potent and
less harmful therapeutics, the optimization of dosing schemes and the development of
suitable routes of administration for the individual patient [698].
10.2. Dynamic Post-Printing Alterations
A challenging aspect of biofabrication is 4D printing, which introduces time as a
fourth dimension and permits materials or components of living cells to alter their shape
or behavior according to stimuli. Tumor modeling can capitalize on this additional mod-
ification after printing, as it can mimic the dynamic characteristics of natural tumors, in-
cluding tumor invasion and ECM restructuring resulting from matrix deposition and en-
zyme-dependent breakdown [699]. The development of fine-tunable biocompatible mate-
rials that react to outside irritants and the response of cells is indispensable for achieving
this. Consequently, stimulus-dependent hydrogels that can remodel and modify their
properties in reaction to external stimuli, such as temperature [700], light [701], moisture
and enzymes, are needed to enable the downstream modification of tumor models.
10.3. Quality Control and Adjustment
The strategy of producing synergistically provides new possibilities for the fast pace,
high throughput generation of advanced cancer models. Regulatory hurdles must be
overcome, though, prior to the full-scale use of these integrative organotypic cancer mod-
els in clinical settings. Additional multidisciplinary work is needed to standardize and
develop guidelines for biological issues, comprising biomaterial and medium prepara-
tion, specimen retrieval and operational guidelines. Real-time readout techniques incor-
porated into organ-on-a-chip systems that utilize machine learning-based image analysis
can be devised to track tumor performance [702]. These techniques could enhance the
process of quality assurance and improve efficiency, uniformity and replicability [703].
While these patient-specific tumor models are impressive with respect to the ethical prin-
ciples of minimizing animal testing, their scaling and clinical translation continue to pose
significant ethical difficulties [704]. Because of the fast pace of technological advances, it
is not straightforward to foresee the potential future applications and banking of these
models, which raises doubts about getting informed permission from donors [704]. Ethical
principles for commercialization are also needed to guarantee a balanced sharing of ben-
efits among donors/patients, scientists, commercial suppliers and other actors participat-
ing in the evolution of these models [1].
10.4. Predictive Integration at the Clinic
Building a 3D tumor model for individual patients in personalized medicine could
be expensive and take time. A potentially more efficacious pathway could be the genera-
tion of large arrays of patient-derived tumor models that reflect certain cancer subtypes,
which could aid in the identification of biomarkers that can predict therapeutic respon-
siveness for a particular patient group. The integration of clinical imaging [705] and ge-
netic profiling [28] can lead to the identification of potential biological or transcriptomic
pathways involved in heterogeneous drug responsiveness. In addition, reliable signatures
of cancer patients that show a response to various therapies can be deduced to make a
forecast for cancer treatment. Multi-organ chip screening of pharmaceuticals relevant to
various co-morbidities is advantageous to enable specific assessment of cytotoxicity and
susceptibility in diverse subpopulations [505]. While the convergence of bioprinting, tu-
mor organoids and organ-on-a-chip techniques is currently in its fledgling stages, a prom-
ising pathway is starting to appear for the creation of integrative organotypic tumor mod-
els for enhanced therapeutic prognostic capability [463,500]. Ultimately, the bioprinting
Cells 2024, 13, 1638 56 of 87
method could surpass animal models and open the door to quicker and more affordable
drug discovery and pathology testing [1]. Nevertheless, multidisciplinary collaboration
between clinicians, biologists, physicists and engineers is necessary to ensure that the bi-
oengineering procedure for cancer models is harmonized, all ethical aspects are resolved
and their use is fully implemented from the laboratory to the patient’s medical care facility
[706].
The future gap of 3D-printed organoids of cancer cells is the localized and dynamic
mechanical analysis of the cells within the organoid in a realistic 3D ECM microenviron-
ment, wherein specific structural or mechanical cues can be altered in a functional man-
ner. Additionally, 3D bioprinting provides the basis for mechanobiological analysis in a
3D environment, whereby cancer cells, neighboring other cell types and the ECM struc-
ture, mechanical properties and composition can be monitored. Consequently, this allows
general mechanisms of cancer progression to be determined based on mechanical and
structural features of the cancer environment in 3D advanced cell culture models.
Table 1. Selected biomedical applications of organoids and assembloids using 3D bioprinting.
Usage
Bioprinting
Technique
Cells
Advantages
Disadvantages/
Limitations
References
Nanotoxicological in-
vestigation
Placement of organ-
oids and assembloids
OBST
Non-small cell
lung cancer line
Calu-3 (Calu-3)
Long-term cultivation of cell
lines based on decreased oxi-
dative stress over time.
Enhanced viability of 3D-
printed cells once incorpo-
rated into the hydrogel.
No loss of time due to cell
passage.
More efficient and cost-sav-
ing.
Consideration of possible im-
pacts on cell morphology and
cell–cell interaction in a 3D
environment.
Dissimilarities between 2D
and 3D data can be uncov-
ered.
The number of cells sown is
strictly determined by the 3D
printing parameters and is
usually lower than the num-
ber of cells imbedded in the
hydrogel-loaded syringe due
to remaining volume in the
Luer lock of the syringe.
Force applied during extru-
sion can mechanically rupture
cells.
[408]
Insertion of physical
limitations in organ-
oids
Physical obsta-
cles
Gastric cancer,
breast cancer,
prostate cancer
and non-small
cell lung cancer
primary cells
Structuring of ECM mechan-
ics in a localized manner and
topographically patterned hy-
drogel scaffolds.
Organoids may create func-
tional vascular systems fol-
lowing transplantation.
Organoids can be utilized to
invert disease-causing muta-
tion to treat disorders caused
by mutations.
Limited cell numbers
Limited resolution of mechan-
ical obstacles
[381]
Hybrid bioprinting
method for
OBB
iPS cell-derived
organoids
Large-scale organoid-like
structures, or assembloids,
Limited size of organoids can
be printed.
[404,707]
Cells 2024, 13, 1638 57 of 87
bioprinting cellular
aggregates (i.e., tis-
sue spheroids and
honeycombs) and or-
ganoids with dimen-
sions between 80 and
800 μm into or onto
hydrogels for both
scaffold-free and
scaffold-based appli-
cations.
can be realized by fast spatial
arrangement.
Heterogeneity of iPS cell-de-
rived organoids is mimicked.
Assembling organoid units of
different shapes pre-divided
into basic organ elements.
Complicated spatial arrange-
ment is difficult due to the
simple form of the block, such
as a sphere.
Changes in bioink characteris-
tics during the printing pro-
cess may hurt organoids.
Spheroids can be
moved through a
shear-thinning hy-
drogel that self-heals
to accommodate the
spheroids and hold
them in 3D space;
also for targeted
merging between
spheroids to form
micro-tissues with
high cell density and
well-defined shape,
which can then be
excised from the hy-
drogel.
Heart disease model
that imitates scar for-
mation after myocar-
dial infarction (MI)
by bioprinting micro-
tissues with spatially
controlled density of
two cell types.
Spheroids are re-
moved from the cell
medium by back
pressure and lifted
into the air to be bi-
oprinted into func-
tional hydrogel, such
as fibrin, collagen,
GelMA or onto sacri-
ficial hydrogels, such
as alginate and aga-
rose.
AAB
Induced pluripo-
tent stem cell
(iPSC)-derived
cardiomyocytes
and primary hu-
man cardiac fi-
broblasts (CFs)
Human MSCs
(hMSCs) were
isolated from
fresh unpro-
cessed bone mar-
row from human
donors
Custom forms for 3D bi-
oprinting have been designed
using computer-aided design
(CAD) software.
Design of personalized in
vitro disease models that de-
liver comparable functional
outcomes to preclinical ani-
mal models while maintain-
ing ease of investigation, min-
imizing expense, and in for-
mats that accommodate a
wide variety of imaging and
assessment techniques.
Facilitate advances in the pro-
duction of assembloids,
where organoids from vari-
ous tissues or tissue regions
are merged in a regulated en-
vironment to investigate tis-
sue development and matura-
tion.
Mimic pathological scarring
characteristics that occur after
myocardial infarction, and
with the help of measure-
ments of cardiac function
(contraction, electrophysio-
logical synchronization),
miRNA therapeutics for repa-
ration could be tested.
Simple model
AAB is poorly suited for gen-
eration of neural assembloids,
as neural organoids display
large diameters, relatively
weak surface tension, and a
propensity to perform plastic
deformation and break down
under relatively low vacuum
force.
[88,407,444,708]
Merging and place-
ment of assembloids
with internal cytoar-
chitecture.
Consists of an iron-
oxide nanoparticle
laden hydrogel and
SPOT
human pluripo-
tent stem cell-de-
rived neural or-
ganoids and pa-
tient-derived gli-
oma organoids
Improves OBB technique
Generates accurately ar-
ranged assembloids com-
posed of human pluripotent
stem cell-derived neural or-
ganoids and patient-derived
glioma organoids.
Magnetic nanoparticle
(MNP)-laden cellulose nano-
fiber (CNF) hydrogel that
may possibly impact cell func-
tion.
[398]
Cells 2024, 13, 1638 58 of 87
magnetized 3D
printer to enable the
regulated lifting,
transport, and depo-
sition of organoids.
Constructs assembloids that
recapitulate major develop-
mental processes and disease
etiologies.
Construction of neural assem-
bloids in 3D.
Construction of assembloids
based on dorsal and ventral
forebrain organoids.
Merging and place-
ment of assembloids
with internal cytoar-
chitecture.
Magnetic levi-
tation
Implementation of internal
structures in organoids.
Mimics the in vivo organs
and tissues more closely.
Relies on cellular internaliza-
tion of a bioinorganic hydro-
gel comprising iron oxide.
Magnetic particles alter cellu-
lar functionality in 3D.
[709]
SPOT = spatially patterned organoid transfer, OBB = organ building blocks, AAB = aspiration-as-
sisted bioprinting.
Funding: Supported by the Open Access Publishing Fund of Leipzig University.
Acknowledgments: I thank Thomas M. L. Mierke for discussion and proofreading.
Conflicts of Interest: The author declares no conflict of interest.
Abbreviations
2PP
two-photon polymerization
3DP
three-dimensional printing with nozzles
AAB
aspiration-assisted bioprinting
AgNPs
silver nanoparticles
AM
additive manufacturing
ASC
adipose-derived stem cells
CAFs
cancer-associated fibroblasts
CHO
Chinese hamster ovary
dECM
decellularized extracellular matrix
DIW
direct ink writing
DLP
direct laser processing
DMD
digital micromirror device
EBM
electron beam melting
ECM
extracellular matrix
ECs
endothelial cells
EHD
electrohydrodynamic printing
EHS
Engelbreth–Holm–Swarm
EY
eosin Y
FDM
fused deposition modeling
FeHA
Fe-doped hydroxyapatite
FRESH
freeform reversible embedding of suspended hydrogels
GAGs
glycosaminoglycans
GBM
glioblastoma
GELGYM
gelatin-based glycidyl methacrylate
GelMa
gelatin methacrylate
HA
hyaluronic acid
HAMA
methacrylated hyaluronic acid
hMSCs
human mesenchymal stem cells
HMVECs
human microvascular endothelial cells
HUVECs
human umbilical vein endothelial cells
Cells 2024, 13, 1638 59 of 87
iPSC
induced pluripotent stem cells
LOM
laminate object manufacturing
LOX
lysyl oxidase
MeAlg
methacrylated alginate
MMP
matrix metalloprotease
MNP
magnetic nanoparticle
MPS
micro physiological system
MSLA
masked SLA
NiPAAm
N-isopropylacrylamide
NorHA
norbornene-functionalized HA
OBB
organ building blocks
PAC
printed active composites
PCL
poly(ε-caprolactone)
PCL-PEG-PCL
polycaprolactone-poly(ethylene glycol)-polycaprolactone
PD-1
anti-programmed cell death protein-1
PD-L1
anti-programmed death ligand 1
PDOs
patient-derived tumor organoids
PEG
polyethylene glycol
PEGDA
poly ethylene glycol diacrylate
PEGMA
PEG dimethacrylate
PEGTA
poly(ethylene glycol/tetracylate
PEO
poly(ethylene oxide)
PLGA
poly(lacte-co-glycolic acid)
PLLA
poly-lactide
pNIPAMA Ac
poly-N-isopropylacrylamide-coacrylic acid
PPy
polypyrrole
PVA
polyvinylalcohol
SA
sodium aliginate
SCE
shape change effect
SLM
selective laser melting
SLS
selective laser sintering
SMA
shape memory alloys
SME
shape memory effect
SMG
shape memory material
SMMs
shape memory materials
SMP
shape memory polymers
SPOT
spatially patterned organoid transfer
TDCs
tissue-derived cells
Tg
transition temperature
TILs
tumor-infiltrated lymphocytes
TME
tumor microenvironment
UAB
ultrasound-assisted bioprinting
UV
ultraviolet
VBP
volumetric bioprinting
μCLIP
micro-continuous liquid interface production
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