Next Evolution in Organ‐Scale Biofabrication: Bioresin Design for Rapid High‐Resolution Vat Polymerization

Abstract

The field of bioprinting has made significant advancements in recent years and allowed for the precise deposition of biomaterials and cells. However, within this field lies a major challenge, which is developing high resolution constructs, with complex architectures. In an effort to overcome these challenges a biofabrication technique known as vat polymerisation is being increasingly investigated due to its high fabrication accuracy and control of resolution (μm scale). Despite the progress made in developing hydrogel precursors for bioprinting techniques, such as extrusion based bioprinting, there is a major lack in developing hydrogel precursor bioresins for vat polymerisation. This is due to the specific unique properties and characteristics required for vat polymerisation, from lithography to the latest volumetric printing. This is of major concern as the shortage of bioresins available has a significant impact on progressing this technology and exploring its full potential, including speed, resolution and scale. Therefore, this review discusses the key requirements that need to be addressed in successfully developing a bioresin. The influence of monomer architecture and bioresin composition on printability is described, along with key fundamental parameters that can be altered to increase printing accuracy. Finally, recent advancements in bioresins are discussed together with future directions. This article is protected by copyright. All rights reserved

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Review
Next Evolution in Organ-Scale Biofabrication: Bioresin
Design for Rapid High-Resolution Vat Polymerization
Caroline A. Murphy, Khoon S. Lim,* and Tim B. F. Woodfield*
C. A. Murphy, K. S. Lim, T. B. F. Woodfield
Christchurch Regenerative Medicine and Tissue Engineering
(CReaTE) Group
Department of Orthopaedic Surgery and Musculoskeletal Medicine
Centre for Bioengineering and Nanomedicine
University of Otago
Christchurch , New Zealand
E-mail: khoon.lim@otago.ac.nz; tim.woodfield@otago.ac.nz
K. S. Lim
Light Activated Biomaterials (LAB) Group
Department of Orthopaedic Surgery and Musculoskeletal Medicine
Centre for Bioengineering and Nanomedicine
University of Otago
Christchurch , New Zealand
The ORCID identification number(s) for the author(s) of this article
can be found under https://doi.org/./adma..
DOI: 10.1002/adma.202107759
techniques. Biofabrication involves the
creation of D scaolds comprised of bio-
materials, bioactive molecules and cells
(see definitions Table1). It has allowed for
the study of cells in a D environment, as
compared to D cell cultures which dis-
play inconsistencies with the in vivo envi-
ronment in terms of morphology, cell–cell
and cell–ECM interaction, proliferation
and dierentiation. D bioprinting, which
is a form of rapid prototyping or solid
freeform fabrication (SFF), is the most
common type of biofabrication technology
employed to create D scaolds. It involves
the selective deposition of cells and mate-
rials in a layer-by-layer approach, through
the controlled assembly of biocompatible
materials, and oers the potential to create
D functional tissues with complex and
hierarchical structures, with extracellular
matrices and cells deposited mimicking the
actual cellular arrangement of the tissue
it is replicating.[–] There are numerous
types of D bioprinting technologies such
as extrusion-based fused deposition mod-
eling and inkjet-based printing.[–] When employing these tech-
nologies, material formulations used are known as bioinks when
comprised of living cells, and biomaterial inks when using cell
free natural or synthetic materials.[] Most often bioinks are com-
prised of hydrogel precursors, which encapsulate the cells to be
printed. These hydrogel precursors form crosslinked networks
upon light exposure or a chemical stimuli postprinting. Although
these techniques oer the controlled deposition of materials it is
still lacking in high resolution and the ability to print complex
architectures representing native tissues. To overcome these
challenges a relatively new research approach for biofabrica-
tion is vat polymerization,[,] which is a form of light assisted
printing.[] Compared with other D bioprinting approaches,
such as extrusion-based bioprinting, vat poly merization technolo-
gies oer significant advantages in printing resolution (down to
the µm scale), complexity, and eciency.[] Vat polymerization
printers have evolved from slow but high-resolution ( µm)
biofabrication technologies to more recent technologies, such
as volumetric printing, which allows rapid biofabrication of cen-
timeter-scale constructs in seconds but at a lower resolution. The
next evolution is achieving both biofabrication speed with resolu-
tion at-scale, as highlighted in the overview in Figure1. However,
this potential for evolutionary advancements, particularly in the
field of tissue engineering and regenerative medicine (TERM), is
restricted by the lack of bioresins available. Vat polymerization
The field of bioprinting has made significant advancements in recent years
and allowed for the precise deposition of biomaterials and cells. However,
within this field lies a major challenge, which is developing high resolu-
tion constructs, with complex architectures. In an eort to overcome these
challenges a biofabrication technique known as vat polymerization is being
increasingly investigated due to its high fabrication accuracy and control of
resolution (µm scale). Despite the progress made in developing hydrogel
precursors for bioprinting techniques, such as extrusion-based bioprinting,
there is a major lack in developing hydrogel precursor bioresins for vat
polymerization. This is due to the specific unique properties and characteris-
tics required for vat polymerization, from lithography to the latest volumetric
printing. This is of major concern as the shortage of bioresins available has a
significant impact on progressing this technology and exploring its full poten-
tial, including speed, resolution, and scale. Therefore, this review discusses
the key requirements that need to be addressed in successfully developing
a bioresin. The influence of monomer architecture and bioresin composi-
tion on printability is described, along with key fundamental parameters that
can be altered to increase printing accuracy. Finally, recent advancements in
bioresins are discussed together with future directions.
1. Introduction
The evolving field of tissue engineering and regenerative medi-
cine is immensely advanced in scaold D biofabrication
©  The Authors. Advanced Materials published by Wiley-VCH GmbH.
This is an open access article under the terms of the Creative Commons
Attribution License, which permits use, distribution and reproduction in
any medium, provided the original work is properly cited.
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still majorly lags behind extrusion-based D bioprinting in terms
of printable biomaterials (including bioinks and biomaterials
inks) available for tissue engineering applications. This distinct
lack of materials is due to the complex nature of the printing
process and the demanding material features required to with-
stand the vat polymerization printing process.[] As biomaterials
suitable for tissue engineering applications typically comprise
of hydrogels, they often lack the mechanical stability to be suc-
cessfully printed. To exploit the full potential of the latest light-
based vat polymerization technologies while advancing the field
of tissue engineering, it is imperative to engineer novel bioresins
capable of being printed while supporting cell proliferation.
Therefore, this review focuses on these key issues to describe key
bioresin design criteria and opportunities to unlock organ-scale
biofabrication, i.e., large centimeter-scale constructs biofabricated
at speed and with high resolution, while supporting cell function
and functional tissue formation.
Furthermore, despite the increasing number of reviews
regarding vat polymerization techniques[,,] there is limited
information that reviews the key fundamentals of developing
bioresins and the processing criteria required for a successful
print. The aim of this review is to introduce the common
forms of vat polymerization technology with a focus on current
research and provide a perspective on the key material prop-
erties. There is a focus on influential printer parameters and
strategies that should be considered when developing bioresins.
Finally, this review also outlines recent advancements in engi-
neering bioresins and highlights future directions.
2. Vat Polymerization Techniques and Principles
2.1. Laser-Based Printing
Although various forms of laser-based printing such as laser-
induced forward transfer,[,] pulsed laser deposition,[,] and
matrix-assisted pulsed laser evaporation[,] have been inves-
tigated for the spatial pattern and proliferation of cells,[,,]
these techniques are used to create thin films or to pattern cells
in D. To create D scaolds for tissue engineering applica-
tions and better to mimic the natural environment and com-
plex architecture in vivo, stereolithography and two-photon
polymerization are commonly used for laser-based printing.
Biofabrication:
2.1.1. Stereolithography (SLA)
SLA was developed by Charles Hull in ,[] it was the first
form of D printing and since then many other forms of addi-
tive manufacturing have been developed such as selective laser
sintering and fused deposition modeling. Similar to other
addictive manufacturing techniques SLA allows for the fabri-
cation of constructs from a computer-aided design (CAD) file,
whereby the architecture of the structure is controlled by the
original CAD design. The fabrication of the D construct by
SLA is based on the spatially controlled solidification of a liquid
resin by photopolymerization. This is achieved by scanning a
Adv. Mater. 2022, 34, 
Figure 1. Overview of the key present and future considerations required to achieve biofabrication of organ-scale constructs at high speed and at high resolu-
tion, though the design and development of novel bioresins for vat polymerization technologies. Vat polymerization techniques incorporating cell-laden or
biocompatible bioresins have evolved from slower but high-resolution biofabrication at the tissue scale (e.g., SLA), to more rapid high-resolution biofabrica-
tion oered by DLP, to latest very rapid biofabrication of larger centimeter-scale constructs in seconds but at lower resolutions (e.g., volumetric printing). The
next evolution in biofabrication is achieving organ-scale (cm) constructs at high printing speed together with high-resolution features and print-fidelity across
various printing platforms. This progression in biofabrication can be unlocked though the development of novel, biocompatible, and bioactive bioresins to
match the rapid advancements made in vat polymerization technologies. The development of these novel bioresins are achievable though understanding the
unique material properties and printer considerations required for vat polymerization. Since many of the material and printer considerations are interrelated,
they all need to be considered collectively in order to achieve future advancements with respect to bioresin design and development.

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laser beam onto a photocurable resin, held in a vat, whereby a
the controlled laser beam follows a predetermined path to cure
the resin to a known depth, causing the first layer to adhere to
a support platform,[] as illustrated in Figure2. After the first
layer is photo-crosslinked the build platform is moved the dis-
tance of one layer depth away from the surface to be recoated
with a fresh layer of liquid resin for photo-crosslinking of the
subsequent layer. This process continues until the D construct
is obtained. The dynamic direction of the laser path is con-
trolled by deflection o a rapidly moving mirror galvanometer
whereby a single photon laser is focused through an objective
lens to activate photopolymerization of the photocurable resin.
Photo-crosslinking is initiated by photons, which deliver the
necessary energy to activate a photosensitive additive called a
photoinitiator. Photopolymerization occurs when the quantity
of photons of adequate energy is absorbed per unit of photo-
sensitive resin.[] This energy creates reactive elements (pre-
dominantly radicals or cations) which induce a chain reaction
causing polymerization of the liquid resin. This transformation
from liquid to solid can be reached when a photon threshold is
exceeded[]
=ao
TI
t ()
where T is the threshold (photons m−), Ia is the light absorp-
tion velocity (photons m− s−), and t is the irradiation time
at the threshold (s). Once the threshold is met the resolution
in the z-direction is dependent on the photon flux (m− s−)
and the irradiation time. The cure depth (layer thickness) is
determined by the energy of the light, which is controlled by
the power of the light source and the scanning speed. The
kinetics of the cure reactions during printing is complex
as it is influenced by various chemical and physical factors
including monomer, initiating system, oxygen present, and
temperature,[] which will be discussed later in the paper.
However, simpler equations are commonly used to describe
the crosslinking kinetic during printing. When using an iso-
tropic and homogeneous medium containing a substance
absorbing at a given wavelength the Beer–Lambert law can be
simplified to give a temporal evaluation of the polymerized
thickness, given as
αα
=



=
1where
0
0
0
e
cLn t
ttT
cF
()
where e is the polymerized thickness, α is the Napierian coe-
cient of molar extinction (l mol− cm−), c it the concentration (mol
l−), t is the irradiation time (s), and F is the incident photon flux
(photons m− s−). Understanding the photopoly merized layer
thickness is essential for high resolution printing. Therefore, con-
trol of light penetration and polymerization depth of the resin are
essential parameters for increasing printing resolution.
With regards the resolution in the x–y plane, this spatial res-
olution is primarily governed by the laser beam spot dimeter.[]
Laser spot sizes are commonly µm[,] however spot sizes
of  µm have been achieved.[] Theoretically the beam spot
diameter is the absolute limiting size of a feature that can
be built. However, due to scanning patterns, the smallest fea-
ture printed is normally . times the laser beam diameter.[]
The width of the cured line depends on the laser beam spot
radius as well as a resin constant as given by the equation[]
Table 1. Definition of terms.
Term Definition
Biofabrication: “the automated generation of biologically functional products with structural organization from living cells, bioactive molecules, biomate-
rials, cell aggregates such as microtissues, or hybrid cell-material constructs, through bioprinting or bioassembly and subsequent tissue
maturation processes.”
Bioprinting: “the use of material transfer processes for patterning and assembling biologically relevant materials—molecules, cells, tissues, and biode-
gradable biomaterials—with a prescribed organization to accomplish one or more biological functions.” 
Bioink: “a formulation of cells suitable for processing by an automated biofabrication technology that may also contain biologically active compo-
nents and biomaterials.” 
Photopolymer resin: a liquid polymer material which undergoes photopolymerization and converted into a solid polymer when exposed to light. Photopolymer
resins are used in lighted-assisted printing technologies.
Bioresin: a hydrogel precursor-based photopolymer resin, which supports cell encapsulation and/or attachment and proliferation. A bioresin is a
specifically engineered bioink for light-assisted or vat polymerization bioprinting techniques, which forms a crosslinked hydrogel network
when exposed to light.
Hydrogel: a material which constitutes a group of polymeric materials. It is a hydrophilic structure, which renders them capable of holding large
amounts of water in their crosslinked D networks.
Light-assisted bioprinting: a bioprinting technique which utilizes light to crosslink a liquid photocurable bioresin in a prescribed architecture to create D solid
constructs.
Vat polymerization: a form of light-assisted bioprinting whereby the liquid bioresin is held in a vat (or tank) during the printing processes. The resin is selectively
cured from which a solid D model is constructed.
Photopolymerization: the process of using light and a photoinitiator to facilitate a photo-crosslinking reaction
Photo-crosslinking: a crosslinking reaction, which is initiated by light to transform a liquid precursor solution into a crosslinked solid material.
Photoinitiator: a photosensitive chemical compound that generates a radical species after absorption of light photons to initiate photo-crosslinking.
Photoabsorber: are dyes or additives that can absorb and attenuate light by competing with the initiator for photons.
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=2
w0
d
p
LW
C
D ()
where Lw is the cured linewidth and W is the laser beam spot
radius, Cd is the cure depth, and Dp is the light penetration
depth.
2.1.2. Two-Photon Polymerization (2PP)
PP is another type of laser-based printing, whereby poly-
merization occurs utilising two photons to excite the added
photoinitiator. This technique has achieved nm resolution
however this comes at the expense of long fabrication times.[]
This method uses focused near-infrared (NIR) femtosecond
laser pulses, with a wavelength of ≈nm, the laser is focused
tightly using a high-numerical-aperture (NA) objective.[] Dif-
ferent to typical SLA one photon absorption, in two photon
absorption, photoinitiators absorb two photons when the ini-
tiator molecules transit from lower energy level to a higher
energy level. Two-photon absorption can be obtained through
sequential and simultaneous absorptions.[] In sequential,
the photoinitiator is excited to a real intermediate state, sub-
sequently, a second photon is absorbed. Due to the presence
of the intermediate energy state the material is absorbing
at a specific wavelength; therefore there will be a surface
eect, thus follow the Beer–Lambert law.[] Whereas more
commonly simultaneous absorption is performed, whereby
there is no real intermediate energy state, rather a virtual
state, where two-photon absorption happens only if another
photon arrives within the virtual state lifetime ≈− s.[,]
As two photons have to be absorbed simultaneously, high
intensities are needed for PP, which can only be provided
by a tightly focused laser beam hence why femtosecond lasers
are used. In order to generate a suciently high density of
radicalized starter molecules the excitation beam intensities
should be of TW per cm.[] Unlike SLA where an initiator
only absorbs one UV photon with a short wavelength through
linear absorption. For PP process, an initiator absorbs two
photons with a long wavelength through nonlinear absorp-
tion[] where the two-photon polymerization rate is propor-
tional to the square of the laser intensity.[] In turn, this leads
to a more localized initiation of the polymerization, and there-
fore higher resolutions.
When using NIR femtosecond laser pulses the liquid resin
is transparent which allows the laser pulses to be focused into
the bulk volume of the resin, rather than the planer surface
of the resin of one-photon SLA.[] Like SLA once a threshold
value of the density of photons is achieved, absorption occurs
within the focal volume, initiating the polymerization process.
The laser is traced throughout the volume of resin creating a
D microstructure. Originally to create D features, the con-
struct was moved during fabrication on a piezoelectric stage,[]
later this was developed to take advantage of a galvanometer
mirror to obtain features in the x–y direction in conjunction
with a vertical axis piezoelectric stage.[] For PP there are
two dierent scanning modes to fabricate the construct, con-
tinuous scanning or pinpoint scanning. The resolution diers
for both types, the fundamental feature size are line width and
voxel size, respectively.[] A voxel is a D point area whereby
the feature size is determined by the lateral and axial dimen-
sions. To predict the voxel size the polymer threshold must
be first be identified. Given a Gaussian laser pulse simplified
equations can be used to determine the voxel diameter (d) and
length (l)[]
,ln
00
20
21/2
dN
tr Nn
C
L
στ
()
=








 ()
σητ
()
=



−








,2 1
0R
20
21/21/2
lN
tz N
C
L ()
Adv. Mater. 2022, 34, 
Figure 2. Scheme of top-down stereolithography (SLA) and bottom-up digital light processing (DLP). Vz represents platform movement in the z-direction.
Reproduced with permission.[] Copyright , Elsevier.

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where r is the focal dimeter, σ is the eective two-photon
cross-section for the generation of radicals, N is the photon
flux constant, n is the number of pulses, τL is the laser pulse
duration, zR is the Rayleigh length, and C is defined as
C= ln[ρ(ρ− ρth)], where ρ is the primary initiator par-
ticle density and ρth is the minimum density of radicals to
allow polymerization (threshold value). When two voxels are
arranged, the overlap δ is an important parameter determining
the surface roughness which is defined as follows[]
dd
d
dxl z
l
δ
=−×− ()
where d and l are the axial and longitudinal dimensions of the
voxel, respectively, and dz and dx are the axial and lateralscan-
ning steps, respectively.
2.2. Projection-Based Printing
Projection-based vat polymerization is similar to laser-based
approach, whereby light is illuminated on a liquid resin to
photo cure each layer in a sequential manner. However, in con-
trast to the laser-based technique, projection crosslinks all areas
of one layer simultaneously. The most common projection-
based printing is digital light processing and continuous liquid
interface production.
2.2.1. Digital Light Processing (DLP)
DLP is an additive manufacturing method that is based on illu-
minating a photocurable resin. DLP is similar to SLA whereby
a vat of resin is used and layers of the construct are fabricated
sequentially, recoating the polymerized layer with a layer of
liquid resin until the full construct is reached. However, the
source of light diers to SLA, DLP technique utilizes a projec-
tion-based technology to initiate photo-crosslinking of the resin
rather than a laser. Each layer of the construct is fabricated by
projecting a D pixelated single pattern of each layer across
the entire platform at once, as illustrated in Figure. This is
done by employing a digital mirror device (DMD), whereby
the material is subject to polymerization and fabrication of the
entire layer at once. DMD is a chip containing up to a million
reflective micromirrors actuated by electrostatic forces. These
micromirrors can individually rotate to ±° with respect to the
surface, which represents mirrors on or o state.[,] When
the micromirror is rotated to −°, the illuminated light is not
reflected into the projection lens but is reflected into a light
absorber and therefore the pixel appears dark, this is classi-
fied as “tilt o” or o.[] When the mirror is rotated to +°
illumination from the light source reflects into the projection
lens as a D pixel, this is classified as “tilt on” or on.[] Sub-
sequently, the light is directed into the resin to cure the resin
solution. This creates a dynamic mask to create a D pattern
on the resin.[] This projection technique using DMDs is also
known as projection stereolithography (PSL) in literature.[] As
the image of each layer is digitally displayed, itis composed of
an array square pixels, this results in each layer formed from
D small rectangular bricks called voxels. As this technique
allows for a complete layer of resin to be cured at once instead
of a single dot in SLA, it results in a significant decrease on
the build time as it is only dependent on the cure depth and
the exposure time, the size in the x–y direction is negligible.[]
Like SLA after the fabrication of a layer, the platform is moved
by a distance that corresponds to the desired layer thickness,
whereby a subsequent layer is fabricated on the previous layer,
thereby the D construct is fabricated by employing a layer-by-
layer approach.
The kinetics during the curing process are highly complex,
however simplified equations are commonly used to describe
the polymerization kinetics during DLP. Most commonly used
is an equation adapted from the Beer–Lambert whereby the
exponential decay of the light intensity as it passes through a
medium in which it is absorbed is described. This semiem-
pirical equation relates the thickness of the crosslinked layer
(cure depth) to the irradiation dose (E (mJ cm−)) as given by
the equation[]
=




ln
dp
c
CD
E
E ()
where Cd is the cure depth, Dp is the penetration depth, E is the
irradiation dose, and Ec is the critical energy. The irradiation
dose applied is a combination of the light intensity (W m−)
and time applied for (s). As can be seen there are three main
parameters that eect the cure depth, E is an operating con-
dition, however, Dp and Ec are the photopolymer resin optical
properties. Dp the depth that the projected light is penetrated
into the resin until a reduction in irradiance of /e is achieved
and Ec is the critical amount of energy required to initiate poly-
merization.[,] By plotting irradiation dose versus cure depth
leads to a logarithmic-linear working curve.[] Ec can be deter-
mined from the intersection of the curve with the x-axis and Dp
can be obtained from the slope of the curve. To allow for photo
curing the irradiation dose (E) must exceed the critical energy
(Ec) to form a solid cured layer. The value of Ec is dependent on
photoinitiator concentrations and oxygen present.[,] There-
fore the energy required to initiate polymerization of the resin
can be altered by increasing time or irradiation does. Further-
more, the cure depth can increase logarithmically with time
and/or applied irradiation dose (E).
2.2.2. Continuous Liquid Interface Production (CLIP)
Recently, Tumbleston etal.[,] developed a novel DLP printing
approach named CLIP. Whereby the recoating process between
layers was eliminated using an oxygen containing “dead zone”
between the vat and the fabricated construct, as shown in
Figure3. This “dead zone” was achieved by using an oxygen
permeable, UV transparent build widow below the vat, which
resulted in an oxygen-containing interfacial resin layer where
free radical photopolymerization is inhibited. As polymeriza-
tion did not occur in the “dead zone” due to oxygen inhibition
it therefore maintained a liquid interface between the vat and
the printed part. This allowed for the continuous drawing to the
part from the vat, causing suction forces to constantly replenish
the resin. The dead zone thickness is dependent on incident
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photon flux (Φ), photoinitiator absorption coecient (αPI) and
resin curing dosage (Dc), given by the equation[]
Dead
zone thickness0
0
0.5
α
=Φ





−
CD
PI
c
()
where C is a constant, Φ is the number of incident photons
at the image plane per area per time, αPI is the product of the
photoinitiator concentration and the wavelength-dependent
absorptivity and Dc is the resin reactivity of the monomer-
photoinitiator. A working curve of the photo-polymerized cured
layer can also be determined using the above along with expo-
sure time (t) and resin absorption coecient (α) give as[]
Cure
dthickness 1ln 0
0
t
D
PI
c
α
α
=Φ




 ()
Taking advantage of this oxygen-inhibited zone, this contin-
uous printing method greatly enhanced print speed times com-
pared to traditional DLP approaches,[] whereby the light expo-
sure, resin renewal, and part movement must be conducted in
separate and discrete steps. Jiang and co-workers have recently
advanced the CLIP process using a novel constrained window
design, namely, Island Window (IW), which allows the for-
mation of an eective oxygen inhibition layer of greater than
µm.[] One of the key factors to the success of continuous
printing is a proper continuous elevation speed,[] which
can allow for print speeds of to mm h− compared to few
millimeters per hour for DLP.[]
2.3. Volumetric Printing
Layer by layer techniques holds intrinsic design limitations
such as overhanging structures need to be supported by struts.
Furthermore, the speed, geometry, and surface quality limita-
tions of additive processes are linked to their reliance on mate-
rial layering.[] To overcome these geometric constraints of
layer by layer light-based techniques, such as DLP and SLA,
multibeam manufacturing techniques have been proposed,
known as volumetric additive manufacturing.[] Shuste and
co-workers were one of the first to report the use of volumetric
printing, whereby complex D structures were fabricated
Adv. Mater. 2022, 34, 
Figure 3. a) Schematic of Continuous liquid interface production (CLIP) printer. Reproduced with permission.[] b) Printed parts using CLIP technique:
i) micropaddles with stems mm in diameter, ii) Eiel Tower model, cm tall, and iii) a shoe cleat >cmin length. Reproduced with permission.[]
Copyright , American Association for the Advancement of Science (AAAS) c) Printed parts by using IW technique: i) Artichoke model,  mmtall,
ii) Turbo shell model,  mmtall, iii) Vase model,  mmtall, and iv) Cage model,  mmdiameter. Reproduced with open access.[] Copyright ,
Elsevier.

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2107759 (7 of 35) ©  The Authors. Advanced Materials published by Wiley-VCH GmbH
using holographic beams.[] This technique removes the time
dependency nature of the layer by layer technique, allowing
for a one-step volumetric print of a full structure in  to
 s.[] More recently, Kelly etal. developed a tomographic vol-
umetric printing method termed computed axial lithography
(CAL), which allows for the creation of arbitrary geometries
volumetrically through photopolymerization.[] This layer-less
fabrication method was inspired by computed tomography
(CT) imaging. Using this technique, the construct is created
simultaneously in a single step by irradiating a rotating vat of
resin photopolymer, from multiple angles with D dynamic
light patterns using a DLP projector, as shown in Figure4.
Each light pattern exposes the whole build volume, the light
dose resulting from a single exposure is insucient to cross-
link the resin. The polymer solidifies only in the selective areas
where the accumulation of multiple angular exposures results
in an absorbed dose overcoming the gelation threshold.[]
This method results in the simultaneous solidification of spe-
cific object voxels, thus allows for the creation of a construct
in one single step rather than sequentially curing layers of
liquid resin.[] The synchronized rotation and patterning are
controlled by a program computed with a Radon transform,
using the reversed principle of CT.[,] Like DLP and SLA
this method also uses models from CAD, however instead of
slicing and building in the x–y direction the construct is built
in the z-direction, as the z-axis is parallel to the axis of rota-
tion of the material volume as shown in Figure . The suc-
cess and resolution of the printed construct is dependent on
the interplay of several physicochemical parameters including
the resin’s viscosity and reactivity, the étendue of the illumi-
nation patterns and the accuracy of the tomographic dose
reconstruction.[]
3. Printer Design Considerations
3.1. Printer Setup
When developing a bioresin for either DLP, SLA or volumetric
applications one must take into account the printer setup as this
has a significant influence on the required material properties.
In vat polymerization DLP and SLA technology two dierent
printer approaches may be used ) top-down or ) bottom-up.
The main dierence between these approaches is the position
of the light source with respect to the vat of resin. The top-
down approach utilizes light positioned above the vat of resin
material. To repeat this curing process and create a D struc-
ture after each consecutive layer the build platform is lowered
into the vat and a fresh layer of resin coats the previous layer
which is subsequently illuminated and cured. The bottom-up
approach is a similar concept, in that the platform moves in
the z-direction, however the light is positioned below the vat
of resin material and the build platform raises above the vat
after each layer is cured as illustrated in Figure. Raising the
build platform allows the surrounding noncured resin to coat
the previous cured layer, and the resin is again illuminated.[]
By contrast, for volumetric printing the light source surrounds
the vat which eliminates the sequential stage movement in the
z-direction.
There are many intrinsic dierences regarding the set-up of
the fabrication process which has a direct eect on the quality
of a printed object. The major dierence between the top-down
and bottom-up is the recoating step of each layer with uncured
resin. With regards this aspect the top-down approach is sim-
pler, given that as the platform is lowered, the resin is allowed
to freely flow over the previously cured layer. However, with
Adv. Mater. 2022, 34, 
Figure 4. Computed axial lithography (CAL) volumetric printing. a) Underlying concept: patterned illumination from many directions delivers a com-
puted D exposure dose to a photopolymer resin. b) Schematic of the CAL system. DLP projector, digital light processor-based projector. c) Sequential
view of the build volume during a CAL print. A D geometry is formed in the material in less than min. d) The D part shown in (c) after rinsing away
uncured material. e) The part from (d), painted for clarity. f) A larger ( mm tall) version of the same geometry. g) Opaque version of the geometry
in (f), using crystal violet dye in the resin. Scale bars: mm. Reproduced with permission.[] Copyright , AAAS.

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2107759 (8 of 35) ©  The Authors. Advanced Materials published by Wiley-VCH GmbH
the bottom approach the cured part will connect to the bottom
surface of the vat as well as the previously cured layer. Conse-
quently, each layer has to be separated from the bottom of the
vat after the illumination and curing of each layer, this is known
as a peeling step. This peeling step is achieved by raising the
platform out of the vat, whereby a pulling force is required to
break the attachment and vacuum between the current layer
and the bottom surface of the vat, this separation process is
illustrated in Figure5. This periodic detachment of the con-
struct from the bottom surface of the vat, introduces stresses
and deformations in the part being printed, which can lead to
several problems.[] For example if the suction force is larger
than the force between the current layer and previous layers
may cause the construct to break and the building process
to fail.[] Pan etal. reported several factors, which play a role
in the final separation force on the construct, which includes
the viscosity of the resin used, the specific geometry and the
layer thickness that is being printed. Additionally the separa-
tion force is reported to increase linearly with the separation
speed.[] Therefore, to reduce the force, this detachment move-
ment performed between each layer increases the overall print
time.
In an eort to reduce the required forces to separate the
cured layer low adhesion materials have been investigated
such as Teflon[] and silicon.[] However, benefits were only
achieved when fabricating parts with a small cross-sectional
area. Dierent mechanics have been investigated to mitigate this
problem such as employing a base rotation mechanism,[,] or
using oxygen to induce a curing inhibitor layer at the vat surface
limiting the separation force required up to %.[,] More com-
monly used is a tilting or shearing motion of the vat to separate
the cured layer from the vat surface.[,–] Zhou et al. intro-
duced a system where a two-channel approach was discussed,
whereby the vat surface was coated by polydimethylsiloxane
(PDMS).[] The PDMS-aided oxygen inhibition at the bottom
surface creating a lubricating layer near the PDMS film, system-
atically the platform was raised a distance of one layer (causing
some elastic defamation of the PDMS film), followed by a
shearing movement of the vat in the x–y plane, which released
the elastic deformation and filled resin between the cured layer
Adv. Mater. 2022, 34, 
Figure 5. a) The separation process in bottom-up printer design: i) layer is cured and sandwiched between the previously cured layers and the vat
surface, ii) separation is initiated at the interface between the last cured layer and vat surface, iii) separation is increasing as the platform moves up
with the cured layers, and iv) physical separation of the cured part from vat surface is completed. Reproduced with permission.[] Copyright ,
Elsevier. b) SEM images of sintered printed lattices strut using i) bottom-up approach, ii) top-down approach (scale bar:  µm),and fracture surface
of iii) bottom-up approach and iv) top-down approach (scale bar:  µm).Reproduced with permission.[] Copyright , Elsevier. c) A failed build
directly after building before removal of the build platform using bottom-up approach. Scale bar: mm. Reproduced with open access.[] Copyright
, Wiley-VCH. d) Images of the sintered crossed layers of parallels (CLP) structure: i) bottom-up approach and ii) top-down approach. Reproduced
with permission.[] Copyright , Elsevier.

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2107759 (9 of 35) ©  The Authors. Advanced Materials published by Wiley-VCH GmbH
and the PDMS film.[,] More recently a commercial business,
Bcreation, patented the use of a shearing mechanism tech-
nology in their printer.[] Another approach used for reducing
the attachment force is employing a tilting motion to aid in part
separation. During separation one side of the vat can be tilted to
allow a slow peel, therefore significantly reducing the required
pulling up force.[,] This tilting technique is used by an indus-
trial bottom-up printer, The Perfactory by EnvisionTec,[] how-
ever, this method is only successful when the parts are located
close to the tilting side. Additionally tilting motions and vertical
movements inherently introduce undesirable deformations
and stresses in the construct under fabrication, including both
compressive and tensile forces.[] To remove the need of extra
motion of tilting or shearing or if parts are too small to ben-
efit from these motions, the direct pulling up approach is also
employed by commercial printers such as the Micro Plus Advan-
tage by EnvisionTec and LumenX by Cellink.
The separation force is a critical reason for the lack of
bioresins available for vat polymerization. Hydrogels used in
tissue engineering applications typically have a low modulus
(< kPa)[,,] to allow for cell proliferation and the diu-
sion of nutrients. Therefore, these materials often lack the
mechanical properties to overcome these separation forces and
present a significant challenge to achieve a consistent and suc-
cessful print. As the separation forces are present in every layer,
if any layer fails, the subsequent layers will be compromised
and will lead to the failure of the final product. In an eort to
overcome these forces other printing considerations need to
be used and investigated to help to obtain an intact structure,
such as overcuring.[] Overcuring increases the green strength
of the construct (discussed in Section.) and can be achieved
by increasing the exposure time of each layer. This phenom-
enon results in an increased bond strength between the layer
being printed and the previous one, which reduces its risk of
breaking. However, using this method, the resolution will
suer and the strength of attachment of part to the vat surface
will also increase.
As top-down is a semicontinuous process, no detachment
step is required at every step, which results in faster fabrication
compared to bottom-up and less forces are applied to the con-
structed part. However, this approach has its own limitations.
First, the vat depth is dependent on the part height, whereby
the part being printed should be completely covered in resin,
therefore there is a need for large quantity of resin to fill the
vat.[] Although the recoating process in bottom-up is a chal-
lenge it also possesses an advantage over top-down, in that the
resin is constrained between the previously cured layer and the
resin tank,[] therefore maintaining the thickness of the curing
layer. Using the top-down approach the thickness of the resin
layer is not homogeneous. As the layer thickness is not as easily
controllable due to the surface tension.[] Defects may appear
during the printing process when a balance between the sur-
face tension values is not achieved, this is due to the intermo-
lecular cohesion forces of the liquid resin are greater than the
combination of the adhesion forces between the liquid resin
and the previous cured layer and between the liquid resin and
external environment.[] This phenomenon can be observed
as a function of increasing contact angle therefore decreasing
wettability. To overcome this, printers may be designed with a
rotating blade or levelling device to aid the spreading of resin
and achieve a more homogeneous layer thickness.[] Using a
blade is particularly advantageous when building a solid struc-
ture however in a porous structure where there is a solid and
liquid phase this mechanism becomes more dicult. Once a
layer is printed and as the blade moves across the part, it will
encounter gaps, which can impose a forced velocity profile
to the resin, leading to resin movement under the blade and
therefore more diculties in homogenously coating the next
layer.[] Furthermore, a challenge with the top-down approach
is its contact with the natural oxygen rich environment, which
inhibits the chain reaction growth by capturing free radicals, so
a higher exposure time is required. Whereas with bottom-up
the liquid resin it sealed from oxygen therefore eliminating the
unwanted oxygen inhibition eect, allowing for faster and more
predictable printing process.[] This oxygen inhibition is fur-
ther discussed in Section..
One of the key advantages of tomographic volumetric printing
compared to SLA and DLP is the elimination of the layer-by-layer
approach, thereby eliminating the need for additional tilting/
rotating mechanisms to aid a recoating process. However volu-
metric printing possesses its own requirements, mainly the need
to minimize the motion of the construct is critical to achieve
desired printing shape and resolution. This can be achieved
using high viscosity resins, which holds the construct stationary
throughout the printing possess.[] As a result, the construct is
freely suspended in the vat of resin which eliminates the need
for the design and fabrication of support struts, giving more
freedom of design compared to top-down/bottom-up approach,
and has allowed for the printing of fluidic ball-and-cage valve
with free-floating elements.[] Additionally, one of the key advan-
tages of volumetric printing compared to SLA/DLP bottom-up
and top-down approaches, is it circumvents some of the chal-
lenges of printing low modulus hydrogels. As the construct
remains stationary during the volumetric printing, minimal
forces are exerted on the construct thus considerably reducing
the risk of force induced failure.[] One of the key considerations
in tomographic volumetric printing is the étendue of the light
illumination source. Light étendue is a critical parameter to be
considered when using volumetric tomography as it defines the
resolution of the print.[] The voxel resolution in the center of
the build volume is determined by the pixel size of the DLP pro-
jector and the magnification of the lens system. Increasing dis-
tance from the center of the build volume, the eective pixel size
increases proportionally to the divergence of the illumination
beam. Therefore the overlap between pixels at the edge of the
build region, leads to the condition that LsNAs= npLvox, where Ls
is the spatial extent and NAs is the numerical aperture of the
illumination source, n is the refractive index of the resin, p is
fraction of overlap, and Lvox voxel resolution in the center of the
build. Therefore to maintain a high resolution both at the center
as well as on the edge of the build volume requires an illumina-
tion source with a low étendue, LsNAs.[]
4. Key Bioresin Properties for Vat Polymerization
The intrinsic properties of the bioresin have a significant eect
on its printability and the overall outcome of the print. When
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developing a bioresin for laser or projection-based vat polymeri-
zation one of the most common factors to consider is the cure
depth, as this determines the layer thickness, and will change
with variations in the composition of the bioresin. Furthermore, it
is essential to know the cure depth to increase the green strength
of the construct in an eort to withstand the printing process.
One of the most integral properties when developing a new
bioresin formulation is its viscosity. However, given vat polymeri-
zation is a relatively new area within the biofabrication and tissue
engineering field there is a lack of data with regards to bioresin
viscosity and its influence on the construct being printed. Herein
we discuss what parameters influence the cure depth, the green
strength, how the viscosity of a resin impacts the printing pro-
cess and considerations that should be made when developing a
bioresin.
4.1. Cure Depth
The cure depth as discussed in Section.., is the depth of each
individual crosslinked layer, which is defined by the supplied light
energy and the optical properties of the resin. The optical proper-
ties of the resin include the critical energy (Ec) energy to initiation
polymerization and the penetration depth (Dp) of the illuminated
light. These optical properties are a direct function of the formu-
lation of the resin,[] with polymer concentration, photoinitiator
concentration and degree of functionalization (DoF) playing a
role in the cure depth.[] Krishnamoorthy et al.[] reported the
higher the concentration the greater number of crosslinking sites
present therefore a greater amount of UV energy is required to
activate crosslinking. Furthermore, as lower concentrations have
greater amounts of uncrosslinked material, allows light to pen-
etrate deeper, compared to higher crosslinked materials, which
tend to absorb more light. An increase in the photoinitiator
concentration results in an increase in Ec, corresponding to a
decrease of the cure depth. The increase of the photoinitiator
concentration can attenuate the incident light radiation, which
can then result in the decrease of the penetration depth. Further-
more, at a higher photoinitiator concentration, more material is
crosslinked near the surface of the solution, which reduces the
UV penetration depth.[,] The DoF also impacts the cure depth,
with lower DoF requiring more energy to result in a similar cure
depth than that with high DoF. Due to the presence of more
functional crosslinking sites on the monomer with higher DoF
is much less resistant to undergo crosslinking and form the
hydrogel compared to monomer with low DoF.[] Although the
working curve is a basic and required concept for investigating
and characterizing new photocurable resins, it must be noted
that is does not provide information regards the mechanical
properties of the cured photopolymer layer or the kinetics of the
polymerization process. It also does not dierentiate between
highly crosslinked rigid solids (commercial resins) and lower
crosslinked softer solids (bioresin hydrogels for tissue engi-
neering application). This becomes a challenge for bioresins as
the cure depth obtained may not provide mechanical properties
to withstand the printing process. Therefore, the working curve
alone is not sucient in determining the printability of such
resins. Most often when developing bioresins a working curve
is obtained followed by investigations of varying the amount of
photoinitiator and light absorber to obtain constructs with
required mechanical properties or green strength.
4.2. Green Strength
The green strength of a printed material refers to the strength
of the construct immediately after printing (before postcuring if
carried out[]). When using more typical hard resin materials,
an increased green strength will allow for ease of handling and
make the construct more mechanically stable to resist deforma-
tion during removal of unpolymerized resin from the scaolds
internal structure.[] Green strength is less of an issue with the
relatively higher molecular weight polymers commonly used
in many commercial rapid prototyping.[] However in mate-
rial research for DLP or SLA applications if the green strength
is not high enough it will cause the construct being fabricated
to delaminate or detach from the build platform ultimately
causing the print to fail.[] When printing soft bioresin hydro-
gels for tissue engineering applications green strength plays a
particular importance in the constructs ability to withstand the
mechanical forces it is subjected to during the fabrication pro-
cess as discussed in Section..
It’s known that the critical energy (Ec) is an important value
to consider when determining print parameters, however the
modulus of a polymer at gel point is normally too low, particu-
larly for hydrogels to survive the build processes. To overcome
this, and provide sucient green strength, Jacobs defines the
excess energy required (Ex) given by[]
=



−




−










ex
p1
1
xc
p
d
d
p
EE
D
C
C
D ()
The excess energy is directly proportional to the critical
energy and inversely proportional to the cure depth, there-
fore the green strength of the construct can be improved by
decreasing the light penetration depth. The light penetra-
tion depth can be influenced by monomer concentration, the
photo-initiator concentration or with the addition of dyes or
light absorbers.[,] Increasing the absorber concentration in
the resin competes with the photoinitiator in absorbing light
thus decreasing the light penetration depth. With the use of an
absorber to achieve a desired thickness the energy required is
also increased.[] This results in a higher crosslinking density
at the surface, overall improving the resolution by allowing
thinner layers with improved strength to be fabricated. Further-
more the addition of a photoabsorber results in an exponential
decay in light intensity, which leads to a gradient in conversion
within each printed layer leading to a variation in the mechanical
properties.[] Therefore with the addition of absorbers the
amount of photoinitiator and irradiation dose used may also
need to be increased to overcome its negative eects.[] One
must also take into consideration if the dye concentration is too
high or the light perpetration is too low, thus a low cure depth,
will impact the sequential layers as they will not be bound to
each other.[] To allow each layer to be bound a certain amount
of over cure is desired. Over cure is the process whereby each
layer is suciently joined the subsequent layer and has been
employed by many researchers to overcome the issue of green
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2107759 (11 of 35) ©  The Authors. Advanced Materials published by Wiley-VCH GmbH
strength and reduce the risk of delamination between layers.
This is achieved by obtaining a layer thickness larger than the
step size of the stage, which causes inter layer binding also
referred to as stitching.[,] This stitching is an integral pro-
cess for the viability of a printed construct, in particular in
the bottom-up approach to overcome the forces when newly
printed layer needs to be detached from the surface of the vat.
It is essential that this separation force does not surpasses the
crosslinked material’s mechanical properties. Overcure may be
achieved by holding the step size constant, while increasing
exposure time and thus, increases the green strength of a D
construct. It must be noted that although a level of overcure
is desirable, if over exposure is too extensive it results in addi-
tional curing into the preceding layer and therefore loss in the
designed features such as pore closure.[] Therefore, when
developing a bioresin, the light penetration, dyes used and pho-
toinitiator must be balanced to ensure a construct with correct
green strength and intact design features is obtained.
4.3. Viscosity
The viscosity of a material is its resistance to flow upon applica-
tion of a stress[] and is one of the most critical parameters in
vat polymerization as it influences polymerization rate, oxygen
inhibition, cure depth, green strength, and the forces applied
to the construct during the peeling step. Moreover, and most
importantly it determines the printability of the resin as the vis-
cosity governs the recoat process between layers. As previously
discussed, between each printed layer, a thin layer of resin must
be coated on top of the last polymerized layer before sequential
photopolymerization can continue. As the flow of the monomer
resin decreases it takes longer to accomplish this step, increasing
the time lag between layers and extending the overall print time.
High viscosity resins may produce bubbles or an insucient
recoat. Therefore, highly viscous polymers should be avoided,
and low viscosity resins are preferred to increase the eciency of
the recoating process. Viscosities suitable for DLP/SLA are typi-
cally Pa s[] but have been reported up to Pa s[] depending
on its molecular weight,[] and viscosities suitable for volumetric
tomographic printing should exceed Pa s.[]
Viscosity of the resin impacts the cure rate as it influences
the mobility of the monomer. Typically, lowering the viscosity of
the resin increases the mobility of the monomers, which allows
the reactive functionalities and monomers to interact faster,
increasing cure rates.[] By increasing the rate of conversion
from a liquid resin to a solid construct leads to decreased overall
construction times.[] However, viscosity of a resin plays an inter-
dependent role on the polymerization rate and oxygen-inhibitory
eect. For example, Studer et al. found that for acrylate-based
resins by increasing temperatures diusion dominates over
reduction in oxygen solubility whereby that at  °C (viscosity of
≈. Pa s) almost no polymerization was observed, however at  °C
(viscosity of ≈ Pa s) crosslinking was increased.[] Typically,
when increasing temperature of a material, oxygen solubility
goes down in most liquid materials; however, when the viscosity
of the liquid is also lowered, diusivity is improved. As the vis-
cosity decreases at high temperature, oxygen diusion into the
resin is accelerated which then amplifies its inhibitory eect.
However in a closed system whereby no oxygen is present, this
inhibitory eect is not seen and crosslinking increases at higher
temperature due to higher molecular mobility which acceler-
ates the polymerization and increases the final conversion.[]
Therefore, not only does the viscosity of the resin during vat poly -
merization needs to be considered for cure rate but also its simul-
taneous eect in regards to oxygen inhibition. Interestingly, in
other words, a lower viscosity can lead to an increase in oxygen
inhibition when oxygen is present therefore reducing cure rate,
which would apply to top-down set ups. However, when no
oxygen is present, for example in a bottom-up set up, a lower
viscosity will allow for increased cure rate.
Viscosity of a resin must also be taken into account with
regards to cure depth and green strength of construct.
Increasing the viscosity results in more material becoming
crosslinked nearer to the surface of the solution. This in turn
reduces the light penetration depth which results in increased
green strength of the construct. However, although using a
higher viscosity resin to increase green strength, increasing vis-
cosity also increases the capillary forces on the construct as it is
lifted from the base of the vat between each layer, increasing the
separation force.[,] These capillary forces can be detrimental
when printing low modulus hydrogels for tissue engineering
and should be given great consideration. As viscosity plays a
role in many factors of the printing process it is imperative
to understand what influences the viscosity when developing
novel bioresins for vat polymerization.
Many factors play a role the viscosity of the resin, and under-
standing how molecular weight and polymer architecture
influence the viscosity is essential when developing bioresins.
Short polymer chains, such as low molecular weight oli-
gomers, and nonlinear polymers including multiarm stars[]
or hyperbranched polymers[] involve fewer interchain entan-
glements, reducing viscosity. For example, - and -arm star
polyisoprenes, with similar molecular weight as linear versions,
display –% lower viscosity.[] Whereby star and hyper-
branched polymer architectures oer an increased number of
reactive sites for crosslinking, aecting cure and crosslinking
rates.[] As star architectures contain multiple polymeric arms
extending from a core makes the nonlinear chains more di-
cult to pack together resulting in fewer entanglements, leading
to lower viscosity and an architecture option suitable for vat
polymerization.[,] Although decreasing the number of arms
results in a decrease in viscosity, this change in viscosity is
however independent of the number of arms if the molecular
weight of the arm becomes too high.[] By taking advantage
of architecture and molecular weight to control the viscosity,
Elomaa etal. successfully printed three-armed photo-crosslink-
able poly(e-caprolactone) (PCL) oligomers using SLA.[]
A strategy to control the viscosity of a resin, rather than its
architecture, is to use two dierent monomers, a high viscosity
monomer and a low viscosity monomer, at various molar
ratios, in order to obtain a resin with a most reactive composi-
tion with desired viscosities. However it must be noted that as
viscosity of the resin may be altered by mixing varying mono-
mers at dierent concentrations, this will also alter the photo-
polymerization kinetics.[,] A more common method to
control the viscosity and polymerization kinetics of solid cure
resins is the incorporation of an additives known as a reactive
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(liquid comonomers) or nonreactive diluents into the monomer
solution. The amount of diluent required is dependent on the
molecular weight of the macromer, with high molecular weights
requiring relatively high amounts of diluent. With regards to
tissue engineering of hard tissue such as bone, high strength
biocompatible and resorbable constructs are desired. These
materials often require high molecular weight monomer and
therefore a diluent to obtain a printable resin. Poly(propylene
fumarate) (PPF) is a common material investigated as a resin
for these applications. PPF linear oligomer has been reported to
have a high viscosity (e.g., for an Mn = . kDa, Đm = . PPF,
the zero-shear viscosity at  °C is  Pa s[]) requiring a
reactive diluent such as diethyl fumarate (DEF) to obtain an
appropriate viscosity of the resin.[–] DEF has been found
to play a dual role when added to PPF, it can act not only to
reduce the resin’s viscosity but also acts as a co-crosslinker, to
amplify crosslinking through the creation of bridging adjacent
chains of PPF.[] Fisher etal. reported the viscosity of the PPF
resin solution decreases more than three orders of magnitude
as DEF content was increased from % to %.[] The sol
fraction was also found to increase when DEF is increased to
%, which correlates to an improvement in the crosslinking
eciency and corroborated by an increase in elastic modulus.
However, increasing the content of DEF from % to %
was found to increase the sol fraction but reduces the elastic
modulus, indicating impaired crosslinking due to the dilution
eect of the DEF. There is a maximum amount of diluent that
may be incorporated into a resin before the network structures
are significantly altered, mechanical properties are reduced, or
D printing of the construct fails.[,] Therefore, it is essen-
tial to find the correct balance of diluent as it can facilitate or
compromise the crosslinking eciency based on the quantity
used. The influence of architecture of PPF on the viscosity
of resins for DLP have also been investigated by Fer at al.[]
Four-arm poly(propylene fumarate) star polymers were shown
to decrease in complex viscosity as the total Mn increased,
allowing printing of PPF with Mn nearly eightfold larger than
the largest linear PPF oligomer printed previously.[] N-vinyl-
-pyrrolidone (NVP) has also been reported as a reactive diluent
to reduce the viscosity of fumaric acid monoethyl ester-func-
tionalized poly(,-lactide) (PDLLA), to be used in tissue engi-
neering. Like DEF, NVP is a reactive diluent as it decreases
the viscosity of the resin while also acting a co-crosslinker,
copolymerizing with fumaric acid and increasing the poly -
merization rates.[] Resins that are functionalized with reactive
groups such as methacryloyl may be diluted with nonreactive
diluents to create biocompatible solid reins using diluents such
as propylene carbonate[,,] or ethyl lactate.[] Photocurable
PDLLA-based resin functionalized with methacryloyl chloride
and applying ethyl lactate as a nonreactive diluent, acts as a
plasticizer for the system. The plasticizer increases the mobility
of macromers and nonreacted end-groups which enables the
crosslinking reaction to proceed to high conversions and high
gel contents.[] Methacrylate-functionalized poly(trimethylene
carbonate) (PTMC-tMA) diluted with propylene carbonate was
found and the higher molecular weights were the most dicult
to process. The networks formed using these high molecular
weights formulations required the largest amounts of diluent
and consequently the lowest crosslinking densities.[] This,
however, leads to fragile, diluent-swollen photo-crosslinked
structures upon photo-crosslinking.[] However, using nonre-
active diluents, shrinkage occurs upon removal of the diluent.
Therefore, this shrinkage must be accounted for in designing
structures.
When printing softer cell laden gel constructs using SLA
or DLP one must take dierent considerations into account.
Although a low viscosity resin is still required if the viscosity
is too low then cell settling may occur, causing an inhomoge-
neous cell distribution within the printed construct.[] The sta-
bility of cell suspension within a bioresin is very important for
the reliable cell distribution post printing, however cells tend
to sediment by gravity in low viscosity materials as illustrated
in Figure6. To address this problem the addition of Percoll
has been reported to overcome the low viscosity resin for SLA
application and to match the buoyant density of the cells to keep
them suspended.[–] Within synthetic polymers the addi-
tion of hyaluronic acid (HA) has been used not only to
inhibit cell settlement during the fabrication process but also
supply a more natural microenvironment for cell prolifera-
tion.[] Furthermore, Lim etal.[] reported the use of gelatin
methacryloyl (GelMA) in a methacrylated poly(vinyl alcohol)
(PVA-MA)-based bioresin, displaying a homogenous cell distri-
bution resulting in bone and cartilage tissue synthesis by the
encapsulated stem cells, as shown in Figure . Recently silk
fibroin (SF) was used as a viscosity modifier, to reduce cell sedi-
mentation in a GelMA-based resin,[] as shown in Figure .
This increase in viscosity may be due to the amphiphilic
characteristics of the SFs enhanced its anity to the GelMA
molecules and formed a sort of bridge between the molecules.
Furthermore, the folded and irregular shape of SF particles
may have induced entanglement of gelatin polymeric chains
increasing the viscosity.
In contrast to DLP and SLA, tomographic volumetric
printing requires high viscosity resins, which in turn benefits
the printing resolution. First, the diusion of reactive species,
such as oxygen, is reduced by using highly viscous resins signif-
icantly reducing gradient-induced oxygen diusion and a poten-
tial blurring of the dose distribution.[] A high viscosity also
counteracts the sedimentation of the construct being printed.
Photopolymerization of a liquid into a solid typically induces a
–% shrinkage, and in tomographic volumetric printing the
increase in density Δρ of the part being solidified can lead to its
sedimentation into the vat of resin. Given a spherical volume,
in a liquid of viscosity µ, a scaling law for the sedimentation
speed v is v/ρ= µ. However by selecting resins with a viscosity
> Pa s, no significant sedimentation was observed over the
short printing time frames ( s).[] Although viscosity plays
a crucial part in the printability of a bioresin, overall limited
studies have reported on the correct viscosity required using
dierent vat polymerization techniques, for a successful print
and a homogenous cell distribution.
5. Key Resin Crosslinking Considerations for Vat
Polymerization
For the development of bioresins for tissue engineering and
regenerative medicine applications, attention has been driven
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toward water-swollen polymer networks known as hydrogels.
The use of hydrogels supports nutrient diusion and can pro-
vide adhesion sites and signalling cues which guide cell growth
and formation of the desired tissue.[] Hydrogels are als o of
keen interest, as they can be modified to display a range of
mechanical, properties which supports the encapsulation and
proliferation of cells. These type of materials have been widely
used in the development of bioinks for extrusion-based applica-
tions.[,–] To form a crosslinked hydrogel network from a
monomer solution, as previously outlined for vat polymeriza-
tion, photo-crosslinking reactions take place at each layer as the
light is illumined on the liquid resin. These crosslinking reac-
tions are dependent on the formulation of the bioresin and the
photoinitiator used and must be taken into account to optimize
the print.
5.1. Photo-Crosslinking Reactions
In vat polymerization, various crosslinking reactions may be
employed to convert liquid resin into a solid construct. The
most common photo-crosslinking reactions used in vat poly-
merization of bioresins include chain growth polymerization
and step-growth polymerization.
5.1.1. Chain Growth Polymerization
Free radical chain-growth polymerization occurs using a photo-
initiator within the resin solution which decomposes upon a
specific wavelength into radicals which serve as kinetic chain
carriers. This radical system includes various stages, such as
Adv. Mater. 2022, 34, 
Figure 6. a) Schematic diagram of DLP D printing, showing i) cell sedimentation with low viscosity hydrogel precursor solution and ii) stable homog-
enous cell suspension in a more viscous solution. Reproduced with permission.[] Copyright , Elsevier. b) Sectional fluorescence images of GelMA
solutions containing cells with varying quantity of silk fibroin. The sedimentation of cells was evaluated by counting the fluorescent cells in each layer.
Reproduced with permission.[] Copyright , Elsevier, ci) Cross-section image of PVMA/GELMA construct, showing the full height displaying no
cell settling (scale bar = µm), and ii) percentage of cells present in each zone, relative to the total cell amount in the whole cross-section, and
iii) alkaline phosphatase staining (red) of MSCs encapsulated in (A) PVA-MA and (B) PVA-MA/Gel-MA after  d (scale bar =  µm). Alizarin red
staining (red) of MSCs encapsulated in (C) PVA-MA and (D) PVA-MA/Gel-MA after  d, (scale bar =  µm) Alcian blue staining (blue) of ACPCs
encapsulated in (E) PVA-MA and (F) PVA-MA/Gel-MA after  d of culture in chondrogenic dierentiation media (scale bar =  µm). Reproduced
with permission.[] Copyright , IOPscience.

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i) initiation, ii) propagation, and iii) termination.[] Radical
generation occurs in this system when the photoinitiator
absorbs light and converts the photolytic energy into the reac-
tive species (radical), initiating photopolymerization. Once
initiation is successful, the generated free radicals react with
specific functional groups along the polymer chain, forming
covalent bonds and reactive radical intermediates. In turn these
reactive radical intermediates react with other reactive groups,
resulting in the propagation phase. The propagation phase is
terminated by a number of dierent mechanisms including
combination of two propagating chains, the free radical is
transferred to another molecule or interaction with impurities
or inhibitors. The rate of photopolymerization can be described
by the following equation[]
/
pp pp 0t
1/23/2
ε
()
[]
=Φ
vk
Ik M ()
where vpp is the rate of photopolymerization, kpp is the photo-
propagation rate constant, φ is quantum yield, ε is extinction
coecient, I is the incident light intensity, kt is the termina-
tion rate constant, and M is the monomer concentration. From
this equation it may be observed that the rate of polymeriza-
tion is related to the monomer concentration by a power of ..
Additionally, the eciency of the photoinitiator with regards
its extinction coecient, influences the polymerization rate
of the resin solution. One major consideration when using a
free radical polymerization system, is that in the presence of
oxygen this polymerization process is hindered by its consump-
tion of free radicals. This oxygen inhibition is very important
in relation to vat polymerization, particularly in the top-down
approach. As the top layer of resin is in the presence of ambient
oxygen which may impede complete crosslinking of each layer
during the fabrication process, which will negatively impact the
entire print. This oxygen inhibition must be taken into account
when developing bioresins, and is discussed in more detail in
Section .. Typically free radical chain-growth is used in vat
polymerization to crosslink methacrylacted monomers, such as
poly(ethylene glycol) diacrylate (PEGDA)[,,] or GelMA.[,]
5.1.2. Step Growth Polymerization
Step growth polymerization also harness light to crosslink
polymer networks in the presence of a free radical initiator
whereby instead of chain growth occurring polymer chains are
predominately crosslinked via a step-growth mechanism.[]
Most common type of step growth polymerization employed
in bioresins is thiol–ene reactions. It must also be noted that
some thiol–ene reactions can occur concurrently with free rad-
ical chain polymerization, which is known as mixed-mode poly-
merization.[] Thiol–ene reactions are highly ecient reactions
of thiols with reactive carbon–carbon double bonds. Thiol–ene
photopolymerizations are based on the radical catalyzed
addition of a thiol to a vinyl functional group. Once a free rad-
ical is generated by either cleaving an initiator which abstracts
the thiol hydrogen or by cleaving the hydrogen directly from
the thiol,[] the thiol–ene reactions are comprised of a two-
step growth mechanism: i) propagation of a thiyl radical
through a ene functional group to form a carbon radical, ii)
followed by chain transfer from the resulting carbon radical to
a thiol functional group, regenerating the thiyl radical.[,]
This reaction takes place in the presence of a crosslinker,
most commonly the reaction of a reactive polymer with a
multifunctional thiolated crosslinker such as dithiothreitol
(DTT) or PEG-SH[,] to form the crosslinked network. The
nature of the thiol–ene mechanism is orthogonal[] such that
one thiol group will react only once with one double bond.
This allows for the formation of homogeneous hydrogel net-
works compare to spatially inhomogeneous networks of free
radical chain polymerization.[] The kinetics of thiol–ene
reaction are determined by the kinetic rate ratio RK= kp/kCT,
where kp is the rate of propagation and kCT is the rate of chain
transfer.[] Whereby when kp is larger the rate is first order
with respect to the thiol concentration, when kCT dominates
the rate is first order with respect to the alkene concentration,
and when kp≈ kCT the rate is half order with respect to both
the thiol and alkene concentration.[] The values of kp and
kCT are dependent on the reaction conditions and the alkene
group used. Therefore the reaction kinetics are dependent on
the reactivity of the alkene used whereby the alkene reactivity
decreases with decreasing electron density of the carbon–
carbon double bond.[,]
Importantly, thiol–ene free-radical addition polymeriza-
tion has several advantages over traditional radical reactions,
including increased material homogeneity and low shrinkage,
which makes it an attractive chemistry to be used in vat poly-
merization[,] and PP lithography applications.[–]
Furthermore due to the nature of interaction between thiol
groups and oxygen, whereby oxygen tends to abstract the
hydrogen from a thiol group to regenerate the thiyl radical
and thus allows continued polymerization renders this reac-
tion not as susceptible to oxygen inhibition compared to tra-
ditional free radical chain growth mechanisms.[,] This is a
positive attribute when considering crosslinking mechanism
to be employed using top-down vat polymerization. Addition-
ally the use of thiol–ene systems are increasing due to their
advantages compared to radical chain-growth polymerizations
involving methacrylate-based formulations including enhanced
biodegradability[] and higher biocompatibility and, therefore
thiol–norbornene has been widely investigated for D printing
and tissue engineering applications.[,,,–] Overall it is
an attractive approach to photo-crosslink hydrogels using vat
polymerization for tissue engineering application.[,]
5.2. Photoinitiators
One of the key elements of successfully printing a construct
using vat polymerization is the choice of photoinitiator used
within the system as it determines the eciency of polymeriza-
tion, therefore influences the printing time and resolution. The
photoinitiator used within the resin should be selected based
on the material being crosslinked, the photoinitiators absorp-
tion of the wavelength of light used within the system and
which crosslinking mechanism is being employed. The most
common photoinitiators for vat polymerization are highlighted
in Figure7. Photoinitiators used for free radical polymerization
are classified as Type I and Type II photoinitiator.[,,] Type I
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photoinitiator system is the most ecient and commonly used
system in vat polymerization. After light absorption, a cleavage
process initiates once the single molecules reach an excited
state, generating free radicals. The initially formed excited sin-
glet state may cleave directly or undergo intersystem crossing
to give an excited triplet.[] Most photoinitiators predominantly
undergo Norrish type I reactions[] whereby cleavage occurs
predominantly at the α-position of the carbon or carbonyl
group (α-cleavage).[] Photoinitiation kinetics can be described
in the following equation:
2
i
i
A
RIfC
Nhv
ε
=
Φ
()
where Ri is the initiation rate, φ is the quantum yield, ε is the
molar extinction coecient, I is the incident light intensity
(units of power/area), f is the photoinitiator eciency, Ci is the
photoinitiator concentration, NA is Avogadro’s number, h is
Planks constant, and v is the frequency of initiating light.[,]
This equation demonstrates that various parameters aect the
performance of an initiator. By increasing the concentration
or increasing incident light intensity will increase the rate of
initiation of the photoinitiator, by transferring more energy to
cause additional cleavage, however detrimental cytotoxic eects
on cells by increasing light intensity and higher initiator con-
centrations must be taken into consideration.[]
In order to determine the most ecient photoinitiator,
the wavelength of the light source within the printing system
must first be taken into consideration together with its molar
extinction coecient, as the incident light required to induce
the cleavage diers in wavelength and intensity.[] The pri-
mary Type I photoinitiators used in biomedical applications
include Irgacure  and lithium phenyl-,,-trimethylb-
enzoylphosphinate (LAP), both may be used with a near-UV
wavelength of nm. LAP has been reported having a much
higher molar extinction coecient compared to Irgacure 
at nm, of  −cm− and  − cm− respectively,[] there-
fore enabling more light to be absorbed at this wavelength,
increasing Ri and therefore the polymerization rate (Rp) as the
initiation rate has an indirect relationship with the polymeriza-
tion rate, given[]
[]
=




2
pp
i
t
1/2
Rk
MR
k ()
where kp is the rate constant for chain propagation, M is
the monomer concentration, and kt is the rate constant for
termination. Additionally LAP also absorbs in the visible range
of nm with a molar extinction coecient of  − cm−.[]
Therefore LAP[–,–] is increasingly used in vat poly-
merization due to its higher molar extinction coecient, faster
polymerization rates, and ability to initiate with visible light[]
compared to Irgacure .[,] LAP is typically used in a range
of .–. wt% for DLP and SLA applications (as per Table2),
however using volumetric tomography printing allows for a
lower photoinitiator concentration of . wt% to print precise
structures, reducing the potential of cytotoxicity risks.[]
Type II photoinitiators are more complex in their initi-
ating system, and are based on a bimolecular reaction. The
two-component system consists of a light absorbing molecule
(uncleavable sensitizer) together with a coinitiator (syner-
gist). After energy absorption, the excited initiator abstracts
a hydrogen atom from a coinitiator, forming H-donor radi-
cals initiating the photo-crosslinking process.[] The most
commonly used uncleavable photoinitiators are camphorqui-
nones (CQ) benzophenones and thioxanthones,[] which
are able to undergo hydrogen-abstraction. Tertiary amines
are the most commonly used Type II coinitiator. Eosin-Y is
a Type II photoinitiator which has been used for tissue engi-
neering, mainly due to its high water solubility.[,] Using
this photoinitiating system eosin-Y extracts hydrogen atoms
from an amine-functionalized coinitiator, such as triethanola-
mine.[,] Figure
Eosin-Y absorbs and reaches an excited state in the visible
light spectrum, and found to have a molar extinction coecient
of   − cm−, at a wavelength of  nm.[] Although
eosin-Y has been used in PP to successfully crosslink HA-Tyr
Adv. Mater. 2022, 34, 
Figure 7. Absorption spectra of most commonly used photoinitiators for vat polymerization.

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Adv. Mater. 2022, 34, 
Table 2. Synthetic and composite resins for biomedical application and bioresin compositions for tissue engineering.
Material Printer
set up
Wavelength Initiator/crosslinker Photoabsorbing
dye/inhibitor
Coating material/
additive
Cells Time point Functional outcomes Feature
size
Ref.
Cells free
PEGDA Smartphone-
enabled DLP
Ru ( × −– × −)/
SPS( × −– × −)
PonceauR
(.–. wt%)
Feature: ≈
µm
[]
PEGDA DLP  nm DPPO Tartrazine,
Methylene
blue, Coccine
 µm[]
Poly(trimethylene
carbonate)
DLP – nm TPO-L
( wt%)
Orasol orange
dye (. wt%)
Propylene
carbonate
Pore:
≈ µm
[]
PEGDA/polydiacety-
lene nanoparticles
LAP (% w/v%) ≈ µm[]
PPF (EnvisionTEC)
DLP
 nm BAPO ( wt%), Irgacure
 (. wt%)
HMB (. wt%) DEF diluent []
PPF (EnvisionTEC)
DLP
 nm BAPO ( wt%), Irgacure
 (. wt%)
HMB (. wt%) DEF diluent Strut:
. µm
Pore:  µm
[]
PPF (EnvisionTEC)
DLP
BAPO
(–. wt%)
TiODEF diluent Post:  µm
Pore:  µm
[]
PEGDA/GelMA DLP  nm LAP
(.–% w/v%)
Acellular
in vivo 
weeks
Wall
thickness:
mm,
Hollow: 
mm
[]
Cells seeded
PEGDA PP  nm Irgacure 
(. w/v%)
Laminin coating NeuroA cell Cell viability and adhesion evaluated.
Functionality showed via formation
of neurofilaments and neuritic extensions.
Pore: µm,
Strand: 
µm
[]
PEGDA DLP  nm LAP ( w/v%) Yellow food-grade
dye
Poly--lysine and
matrigel
T fibroblasts Day  Showed adhesion and proliferation of fibroblasts
seeded onto printed scaold, as well
as subsequent formation of fibrous-like
cell layer. Functionality of printed
cells not evaluated.
 µm[]
PEGDA DOPsL  nm Irgacure  (.% w/v) BT breast cancer
cells and induced
pluripotent stem cells
(iPSCs)
Day  Functionality assessed for embryoid bodies
formed by IPSCs. Presence of cavities and
the three germ layers observed after  days
of culture using immunohistostaining.
Well:
 µm
[]

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Material Printer
set up
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dye/inhibitor
Coating material/
additive
Cells Time point Functional outcomes Feature
size
Ref.
PEGDA DOPsL  nm LAP (% w/v%) HMBS and TEMPO
(.% w/v)
Adipose-derived stem
cells (ADSCs)
Day  Dierentiation capability of ADSCs seeded
onto the printed hydrogels dier in response to
dierent patterns. Both symmetric fork and asym-
metric fork patterns induced chondrogenic and
osteogenic dierentiation of ADSCs. By contrast,
stripes patterns induced uniaxial ADSCs align-
ment, and supported smooth muscle cell lineage
progression.
 µm[]
PEGDA  nm Irgacure  ( wt%) Tinuvin  (.
wt%)
Fibronectin A bone marrow pro-
genitor cell line
Day  Only showed cell adhesion onto printed
construct with limited functionality
assessment.
Pore:
µm,
Wall:  µm
[]
PEGDA DOPsL – nm Irgacure  (.% w/v),
Irgacure  (.% w/v)
TEMPO (.–
.% w/v), HMBS
(.% w/v)
Fibronectin Human mesenchymal
stem cells (hMSC’s)
and murine myoblasts
CC
Day  Only showed cell morphology after seeded
onto printed construct. Cell functionality
not examined.
≈ µm[]
PCL DLP Blue light Irgacure  ( wt%) Orasol Orange
G dye (.wt%)
Vitamin E
(.wt%)
NIHT fibroblasts Day  Adhesion and proliferation of cells seeded
onto printed scaolds characterized.
Pore:
 µm
[]
PEGDA/PEG DOPsL  nm Riboflavin or fluorescein
and TEA as co initiator
and DPI as catalyst
Human lung adeno-
carcinomaderived
A cells
Day  Limited cell studies, only showed metabolic
activity as an indication of scaold toxicity
≈ µm[]
PEGDA or PEGDMA SLA  nm Irgacure  (. wt%) FITC-–dextran,
PEG–RGDS or
PEG–RGDS–FITC
Human dermal
fibroblasts
Day  Showed cell adhesion and proliferation
via imaging.
 µm[]
PDLLA-PEG-PDLLA DLP  nm Lucirin TPO-L (. wt%) Phenol red
(.wt.%), Hydro-
quinone (.wt.%)
Human mesenchymal
stem cells (hMSC)
Day  Cell adhesion and proliferation showed via light
microscopy imaging and scanning electron
microscopy.
Pore:
 µm
[]
PEGDA or HDDA
or BEDA
DLP  nm BAPO (.wt%-wt%) Lung cancer epithelial
cells A
Day  Cell viability assessed using live/dead staining []
PTMC Envisiontec  nm Lucirin TPO-L (. wt%) Orasol Orange G
(. wt%)
HUVECs Day  Cell proliferation observed using light
microscopy imaging.
Wall:
 µm
[]
PTMC Envisiontec  nm Lucirin TPO-L ( wt%) Orasol Orange G
(. wt%)
Chondrocytes Week  Chondrocytes were able to attach on the
printed constructs and secreted sulfated
glycosaminoglycan (cartilage-specific proteogly-
cans) over  weeks culture period.
Pore:
 µm
[]
PDLLA Envisiontec  nm Irgacure  (. wt%),
Lucirin TPO-L ( wt%)
Orasol Orange G
(. wt%), hydro-
quinone inhibitor
(. wt%)
Ethyl lactate
(%),
Mouse pre-osteo-
blasts
(MCT cell line)
Day  Cell adhesion and proliferation observed
via light microscopy, tissue formation
not evaluated.
Pore:
 µm
[]
PPF DLP  nm BAPO ( wt%) TiO ( wt%) DEF diluent Human mesenchymal
stem cells (hMSCs)
Day  Cell adhesion and matrix production visualized
using scanning electron microscopy
Pore:
 µm
[]
Table 2. Continued.

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Table 2. Continued.
Material Printer
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additive
Cells Time point Functional outcomes Feature
size
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PPF SLA  nm BAPO (– . wt%) DEF diluent,
RGD, cyclo RGD,
and an RGD-
KRSR mixture
MCT-E
pre-osteoblasts
Day  Cell proliferation and osteogenic dierentiation
evaluated. Alkaline phosphatase activity
measured as an indication of degree of
osteogenic dierentiation.
Strand:
 µm
Pore:  µm
[]
PPF/BMP- loaded
PLGA microspheres
SLA  nm BAPO ( wt%) DEF diluent MCT-E
preosteoblasts
Week  In vitro osteogenic dierentiation of attached
cells evaluated using RT-PCR. Gene expression
of alkaline phosphatase, osteocalcin and
collagen type I was measured. Acellular printed
constructs containing BMP- were implanted
into the cranial defect. Bone formation is
measured as reduction in defect area.
Strand:
 µm
[]
PPF/BMP- loaded
PLGA microspheres
SLA . nm BAPO ( wt%) DEF diluent,
apatite, RGD
MCT-E
preosteoblasts
Day  Histological staining (haemotoxylin and eosin)
was used to study cellular adhesion
on printed constructs.
Strand:
 µm
Pore  µm
[]
PEGDA/chitosan SLA  nm Irgacure  (. w/v%) Human mesenchymal
stem cells (hMSCs)
Day  Cell adhesion, proliferation
and coverage area evaluated.
[]
PEGDA/GelMA DOPsL Irgacure  (% w/v) TINUVIN 
(.% w/v), HMBS
(.% w/v) and
TEMPO (.% w/v)
RGDS HUVECs and
NIH-T
Day  Cell morphology was evaluated in response
to printed hemisphere structures using
immunofluorescence staining.
Well:
 µm
[]
PEGDA/BSAMA SLA  nm Ru (.wt%)/SPS
(.wt%)
NIH/T murine
fibroblasts
Day  Viability of cells attached on the
printed constructs evaluated
using live/dead staining.
Strand:
µm Pore:
 µm
[]
Cell encapsulated
PEGDA PSL/DLP EnvisionTec
(Perfactory)
LAP (.–.% w/v%) Percoll Human adipose-
derived stem
cells (hADSCs)
Day  Cell viability and metabolic activity
evaluated, hADSCs dierentiation
was not examined.
Pore:  µm[]
PEGDA SLA  nm Irgacure  and HMPP
(.–.% w/v)
RGDS Human dermal
fibroblasts (HDFs)
Day  Cell viability evaluated to be
≈% after  day post-fabrication.
Hollow: 
µm
[]
PEGDA SLA Irgacure  (.% w/v) RGDS NIH/T cells Day  Long term cell viability, proliferation and
spreading within the multi-layered
printed structures studied.
[]
PEGDA DLBP/SLA  nm VA- (% w/v) Gelatin coating MCF- cells Day  Only cell viability in short-term culture
showed. Long-term functionality
assessment not reported.
Strand: 
µm
[]
PEGDA µSL Irgacure  (.% w/v) Fibronectin Murine OP- marrow
stromal cells and
Mesenchymal
stem cell (MSC)
Day  Osteogenic dierentiation of encapsulated
cells evaluated using brightfield microscopy
and histology staining (Von Kossa staining
to show osteogenic dierentiation).
Strand:
µm,
Pore:  µm
[]

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dye/inhibitor
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additive
Cells Time point Functional outcomes Feature
size
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PEG-co-PDP DLP LAP (.% w/v) Percoll, RGDS HUVECs Day  Viability and metabolic activity of HUVECs encap-
sulated within the hydrogels evaluated.
- []
PEG-PQ PP  nm DMAP (.%) RGDS HUVECs Day  Cell organization within the hydrogel evaluated
using fluorescence imaging.
 µm[]
PEGDA/GelMA PSL Eosin Y NIH T fibroblasts Day  Only cell viability measured within  days culture
period, no other cell functionality assessed.
≈ µm[]
PEGDA or PEGDA/
GelMA
DLP LAP (× − ) Xanthan gum Human Red Blood
Cells, endothelial cells
(HUVECs), primary
human mesenchymal
stem cells (hMSCs),
HEK hepatocytes
Day  hMSCs can be printed and undergo osteogenic
dierentiation. Hepatocytes aggregates were
printed along with HUVECs, and showed
improved tissue engraftment. Functionality
of the liver tissues evaluated by transplantation
into a rodent model of chronic liver injury.
The encapsulated hepatocytes exhibited
albumin promoter activity after  days
of engraftment.
≈ µm[]
PEGDA/PCL/
Collagen
DLP LAP (. w/v%) Milk powder (.%)
and orange food
color (.%)
Percoll CHT/ mouse
mesenchymal stem
cells (ATCC)/human
umbilical vein endo-
thelial cells (HUVECs)
Day  Only viability presented after printing. ≈ µm[]
PVAMA/GELMA DLP Ru (. × −)/SPS(×
−)
PonceauR( wt%) Humanmesenchymal
cells(MSCs), articular
cartilage derived pro-
genitor cells (ACPCs)
Day  Encapsulated stem cells or progenitor cells able
to dierentiate into tissue of interest. Immuno-
histological staining used to evaluate extracellular
matrix formation within the printed constructs.
- µm[]
PEGDA/GelMA DMD-PP Irgacure  (.% w/v) NIH-T fibroblasts Day  Cell viability and morphology evaluated.  µm[]
-PEGDA DLP  nm LAP, TPO, Irgacure ,
Irgacure , Eosin-Y,
Riboflavin, Omnirad TPO-L,
Irgacure , Irgacure -
DW, VA-, Irgacure 
Maxgard R,
Maxgard R,
TEMPO, HMBS,
Hydroquinone
Mouse myoblast
CC
Day  Cell viability and metabolic activity evaluated.
Histology staining (hemotoxylin and eosin)
used to visualize cell distribution within printed
construct.
≈ µm[]
PEGDA: poly(ethylene glycol) diacrylate; PEGDMA:poly(ethylene glycol) dimethacrylate; PCL: polycaprolactone; PTMC: poly(trimethylene carbonate); PDLLA: poly(,-lactide) acid; HDDA: ,-hexanediol diacrylate; BEDA:
bisphenol A ethoxylate diacrylate; PPF: poly(propylene fumarate); PTMC: poly(trimethylene carbonate); HDDA: ,-hexanediol diacrylate; BEDA: bisphenol A ethoxylate diacrylate; BSAMA: bovine serum albumin.
DOPsL: dynamic optical projection stereolithography; DMD-PP: digital micromirror device projection printing; PSL: projection stereolithography; DLBP: direct laser bioprinting.
LAP: lithium phenyl-,,-Trimethylbenzoylphosphinate; DPPO: diphenyl (,,-trimethylbenzoyl) phosphine oxide; TEA: triethanolamine; DPI: diphenyleneiodonium chloride; BAPO: phenyl bis(,,-trimethylbenzoyl)-phos-
phine oxide; RU: ruthemium; SPS: sodium persulfate; HMPP: -hydroxy--methyl--phenyl--propanone); DMAP: ,-dimethoxy--phenyl- acetophenone.
HMBS: -hydroxy--methoxy-benzophenon--sulfonic acid; TEMPO: ,,,-tetramethylpiperidine-oxyl.
FITC: fluorescein isothiocyanate; DEF: diethyl fumarate, HMB: -hydroxy--methoxybenzophenone; KRSR: lysine–arginine–serine–arginine, TiO: titanium dioxide.
Table 2. Continued.

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hydrogels through dityramine crosslinking,[] and in SLA to
successfully crosslink PEGDA and GelMA,[] it is not widely
used in vat polymerization due to its need of coinitiators and
accelerants to speed up the reaction. Another Type II photo-
initiating system, which has been used in recent years is
tris(,-bipyridyl) dichlororuthenium (II) hexahydrate (Ru),
with a coinitiator of sodium persulfate (SPS). Ru is based on
a transition metal complex and has previously reported as
being highly absorptive in the visible light range, with a molar
extinction coecient of  –  − cm−.[–] The Ru+
metal complex is known to produce an excited state capable of
donating an electron to persulfate resulting in cleavage of the
OO bond resulting in Ru+.[,] Once SPS accepts an elec-
tron it dissociates into sulfate anions and sulfate radicals. The
newly generated sulfate radicals are able to either trigger free
radical chain polymerization or thiol–ene photo-crosslinking.
Ru and SPS have been used to crosslink tyramine functional-
ized synthetic polymers and tyrosine-rich proteins[] and has
been used in DLP applications.[,,]
5.3. Oxygen-Inhibited Free Radical Photopolymerization
As previously discussed oxygen inhibition may be advan-
tageously used in bottom-up vat polymerization when an
oxygen-permeation element is employed to alleviate forces
experienced by the printed part during the peeling process.[]
With regards to tomographic volumetric printing, the presence
of oxygen could potentially result in a gradient-induced oxygen
diusion, which could cause blurring of the dose distribu-
tion over time. However, given more viscous resins are used
(> Pa s) the resulting diusion of oxygen radical scavenges
is less than µm over the short manufacturing times required
( s). Therefore diusion blurring of the dose distribution is
deemed negligible for this approach.[] However, for top-down
vat polymerization, oxygen inhibition is one of its major draw-
backs, due to its inhibiting eect on the photopolymerization
reaction at the surface of the liquid resin. Therefore, this issue
needs to be understood and accounted for when using this
process.
Polymerization can occur through various photo-crosslinking
reactions such as free radical chain polymerization and thiol–
ene reactions as previously discussed. Oxygen inhibits free
radical polymerization however thiol–ene reactions are insensi-
tive to oxygen.[,,] During free radical polymerization, the
generated free radicals can react with specific functional groups
on the polymer chains to form crosslinks within the material.
However in the presence of oxygen, free radical polymerization
is inhibited by oxygen either due to quenching the excited state
photoinitiator or by forming a peroxide upon interaction with a
free radical.[,,] Therefore, using top-down technique may
cause unpredictable crosslinking due to the material in contact
to the natural environment containing oxygen. Many strategies
have been employed to reduce this oxygen inhibition[] and
should be given great consideration when developing bioresins
for top-down vat polymerization. Some strategies to prevent
oxygen inhibition during photopolymerization includes phys-
ical barriers such as the addition of wax barrier coat to slow
down the diusion of oxygen[] or curing in an oxygen free
environment by covering the resin being polymerized with an
inert gas, typically nitrogen, helium[] or carbon dioxide.[]
Oxygen inhibition in acrylate-based resins may be altered by
taking the permeability and solubility of oxygen into account,
through changing the temperature of the resin and thus the
viscosity of the resin as previously discussed.[] Another pos-
sible solution to overcome oxygen inhibition is the use of addi-
tives, such as such as tertiary amines, however these are not
appropriate in all cases whereby for mixed epoxy/acrylate resins
they retard cationic polymerization.[]
The aforementioned strategies may be dicult to employ
when using cell laden materials, therefore the simplest method
of mitigating oxygen inhibition is to increase light intensity or
photoinitiator concentration, thus increasing number of radi-
cals formed when exposed to light. This increase in radicals
can rapidly combine with oxygen lowering its overall negative
eect. Increasing the light intensity is more eective in miti-
gating oxygen inhibition of acrylate-based resins compared to
increasing radiation dose. Whereby it has been reported that
curing for a short time with a high intensity lamp provided a
% double-bond conversion at the surface, while curing with
the same total radiation dose using a longer time frame and a
lamp with three orders of magnitude lower intensity gave only
% double-bond conversion.[] Although light intensity can
show benefits in synthetic materials when increasing intensity
for cell laden bioresin one must take into account cell viability
when increasing light intensity and radicals formed. Lim etal.
reported using Irgacure  to crosslink GelMA, increasing
its concentration from . to . wt% or light intensity from
 to mW cm− to overcome oxygen inhibition significantly
reduced the cell viability. However, diering the photoinitiator
system and employing Ru/SPS and visible light system resulted
in higher cell viability.[]
6. Key Considerations of Additives Required
for Vat Polymerization
When developing a bioresin formulation for vat polymeriza-
tion, one should consider the use of a dye or pigment to create
high resolution features such as complex architectures or per-
fusable structures. These dyes are typically photoabsorbers,
which can be used to attenuate light and increase the resolution
of the print.
6.1. Photoabsorbing Dyes and Pigments
Photoabsorbers and photoinhibitors are widely used in DLP
to improve the printability of intricate and functional archi-
tectures.[,] Photoinhibitors such as TEMPO act to mitigate
free radical migration distance therefore reducing structure
bleeding while creating higher aspect ratio structures.[,]
However more commonly a photoabsorbing species is added
to the resin as they function to attenuate light by competing
with the initiator for photons. By tuning the amount of light
absorber used, the layer thickness can be controlled as the
photo absorber results in a decrease in penetration of the light.
It is important that the limiting of light penetration is not so
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great that the initiator cannot function. The function of the
photoabsorber can be understood through Beer–Lambert law,
assuming that the monomer has an absorption constant α and
a light intensity Io. the intensity become I at depth z to give the
following equation[]
α
()
=−
exp
0
II
z ()
Given I(Cd) = Icure, where Icure is threshold intensity at which
the monomer cure. When a photoabsorber is added, to the
resin α becomes larger (α′) and CD becomes smaller (CD′) the
addition of the absorber to the monomer increases the absorp-
tion constant and decreases the curing depth, and thus the
structural resolution is therefore enhanced.
Commonly investigated photoabsorbers include natural or syn-
thetic food dyes, which are fully dissolved in the resin. Tartrazine,
a yellow food dye colouring, has an absorbance peak at nm,
is a promising photoabsorber from bioprinting applications
due to it biocompatibility, low toxicity, and wide use in the food
industry. Grigoryan etal. investigated tartrazine to create internal
vessel formation and the eects of photoabsorbers on the gela-
tion kinetics of photopolymerized hydrogels. For photopolymeri-
zation characterization with short duration light exposure, indi-
cated that photoabsorbers cause a dose dependent delay in the
induction of photo-crosslinking.[] The group demonstrated that
the addition of tartrazine in PEGDA hydrogels to create complex
multivascular networks and show the feasibility of intravascular
oxygen transport between D entangled networks. In addition,
the use of the photoabsorber allowed the group to fabricate an
alveolar model with perfusable open channels.[] More recently,
Yang et al.[] also investigated PEGDA with various photoab-
sorbers including methylene blue, coccine, and tartrazine. It was
found that at a wavelength of nm, tartrazine had the highest
absorbance and was more eective in slowing down the light-
induced crosslinking reaction compared to methylene blue and
coccine. Furthermore, the photoabsorber was able to minimize
light scattering into the designed hollow channel thereby slowing
down over curing and facilitating the printability of internal struc-
tures. Tartrazine has been found to have an optimal concentra-
tion, and therefore has a limited eect on the printability. Other
photoabsorbers which may be used include curcumin, derived
from turmeric (absorbance peak at ≈ nm)howevergiven its
lipophilic it does not wash out form the structure postprinting.
Anthocyanin, derived from blueberries, given its absorbance
peak at ≈ nm require much higher concentrations than tar-
trazine to provide suitable light attenuation under visible light.[]
Ponceau R (red food dye) was studied by Lim etal.[] demon-
strated that addition of  wt% of the absorber was suitable for
a poly(vinyl alcohol)-methacrylate (PVA-MA) resin, allowing
improved control over print resolution. Other widely used photo-
absorbers are -hydroxy--methoxybenzophenone--sulfonic acid
(HMBS)[,,] and TINUVIN .[] HMBS is an attractive
photoabsorber used in biomedical applications as it is an FDA
approved chemical used in sunscreen and cosmetic products.[]
Inorganic particles and pigments may also be used in vat
polymerization to achieve complex structures by absorbing
excess light and confining the polymerization area. Gold
nanoparticles may be used due to their biocompatibility and
light attenuating properties however they become physically
entrapped and make transmission or fluorescence micros-
copy impractical on the constructs.[] Pigments are substances
that accomplish the same thing as dye photoabsorbers but
does not dissolve. A common pigment researched is titanium
dioxide (TiO) as it has been found to be biocompatible with
a high absorption potential however they also have the unde-
sirable eect of scattering light.[,] This light scattering can
promote crosslinking in areas with no light projection known
as “dark cure,” which results in loss of lateral resolution and
accuracy.[,] To overcome this issue dyes and pigments may
be used in conjunction with each other to achieve optimum
results. Mott et al.[] study the use of TiO and -hydroxy-
-methoxybenzophenone (HMB), commercially known as oxy-
benzone, as a coattenuate. TiO contributed to reducing the
cure depth, where HMB performed a dual purpose of reducing
the cure depth while simultaneously absorbing the scattered
light of TiO, thus eliminating “dark cure” regions.
An extensive list of UV blockers used in sunscreen products
may be employed by vat polymerization, such as UVA (benzo-
phenones, anthranilates, and dibenzoylmethanes) and UVB
filters (PABA derivatives, salicylates, cinnamates, and camphor
derivatives). However, due to the technologies relative new use
compared to sunscreen, only a low range of dyes and pigments
have yet been investigated with vat polymerization.[]This
highlights the potential for a number of additives to be investi-
gated for use in developing bioresins.
7. Bioresins and Recent Advances
A major focus for vat polymerization for biomedical appli-
cations is the development of bioresin to be used with the
technology. This is a major challenge within the research com-
munity due to the complexity of the system and the consid-
erations that need to be accounted for as discussed throughout
this review. Despite these challenges many researchers have
developed bioresins, however the use of synthetic resins has
predominantly been explored. Synthetic materials oer the
wideset range of tailorable chemically and physically proper-
ties.[] In particular, PEGDA has been extensively used due
to its mechanical properties and ease of manufacture. As nat-
ural-based monomers have inherited variability this can lead
to more unpredictable printing outcomes, which have limited
their use in light-assisted biofabrication. Furthermore, a major
challenge lies in developing soft bioresins for cell encapsula-
tion that can withstand the conditions of the printing process,
which hinders the use of vat polymerization for biofabrication
applications.[] Therefore, although the library of photopolym-
erizable bioresins continues to expand it is still majorly lacking
in naturally derived monomer for the development of high
resolution prints including encapsulated cells. The following
section outlines the common resins used for biomedical appli-
cations and highlight the recent advances made in the field.
7.1. Synthetic-Based Bioresin
As previously mentioned, the most widely used materials are
synthetic-based materials due to their ease of fabrication, with
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PEGDA being the most predominant material used in vat
polymerization. PEGDA is a synthetic non-degrading water con-
taining hydrogel.[] It is used in biomedical applications due to
its biocompatibility and tuneability. For instance although PEG
is non-degrading, hydrolytically degradable PEGDA hydrogels
have been developed by utilising a step growth polymeriza-
tion reaction between PEGDA and DTT to achieve a predict-
able degradation rate.[] Like many monomers PEGDA can
be synthesized and used at a variety of molecular weights with
common range of [] to  Da[,] used in literature.
Recently, Grigoryan etal. used PEGDA (Mw = Da) to pre-
pare complex volumetric structures with DLP printing, which
successfully modeled a noncellular vascularized alveolar unit
(Figure8).[] Although PEG displays many positive attributes,
its poor adhesion for cells retards its use in biomedical applica-
tions. For cell attachment PEG scaolds still require grafting
of peptide sequences such as RGDS[,] or adhesive proteins
such as fibronectin[] or laminin.[] For example, Zhang
et al. fabricated various microarchitectures and illustrated
that PEGDA functionalized with RGDS were capable of sup-
porting cells adhesion and proliferation.[] Later PGEDA was
functionalized by coupling fibronectin to the hydrogel surface
to promote cell adhesion. The group printed fractal forms of
biological looking patterns to mimic the native bifurcating mor-
phology of branching biological structures. Both murine myo-
blast cells and human mesenchymal stem cells were found to
adhere and grow on the fractal geometry scaolds (Figure).[]
Others have investigated the combination of synthetic
materials in conjunction with natural materials to create a
bioresin that can be printed at high resolution, which also sup-
ports cell encapsulation and proliferation. For example, Lim
et al. investigated the use of PVA-MA/GelMA composite as a
bioresin for DLP toward the development of free form living
tissue analogues.[] GelMA is known to support cell adhe-
sion, growth and proliferation, whereby its addition to PVAMA
would result in a biosynthetic hydrogel resin, possessing bio-
functionality. The incorporation of GelMA to a PVA hydrogel
led to a significant increase in the number of cells attached and
was crucial for long term cell viability. The bioresin showed
capabilities of supporting both bone and cartilage tissue for-
mation. Smith et al.[] also recently investigated the use of
composite resins comprising of % methacrylated bovine
serum albumin (BSAMA) with varying levels of PEGDA. It
was found that BSAMA alone was not printable however by
harnessing the benefits of PEGDA and its addition to BSAMA
made the resin printable. Excellent cell viability was found after
seeding constructs and culturing for  days, however viability
was found to decrease with increasing PEGDA concentration.
Others have used PEGDA as a material to create cell encapsu-
lated gradient features[] and others explored its use in devel-
oping scaolds with regionally varied stiness,[] bio-inspired
detoxification modules,[] peripheral nerve guidance con-
duits.[] Along with PEGDA other synthetic materials used
for bioresins include PDLLA-PEG-PDLLA, poly(propylene
fumarate) and PEG-peptides.[–] An overview of synthetic
and composite bioresin formulations in literature are given in
Table. The use of proteins and natural-based materials in con-
junction with synthetic materials have led to an increase in the
development of synthetic bioresins. However as evident from
Table , there is a distinct lack of studies reporting encapsu-
lated cells, with a predominant focus on cell seeding and inves-
tigation at early time points. For tissue engineering applications
it is essential to develop functional bioresins that promote cell
proliferation and support ECM production to regenerate dam-
aged tissues. Thus, there is a critical need to develop bioresins
which prove to support cells growth over extended periods and
investigate the synthesis of newly formed ECM.
7.2. Natural-Based Bioresin
Limited studies have been conducted on fully natural-based
bioresins. The key advantage of using naturally derived mate-
rials is the presence of biologically active components that
allow cells to adhere and migrate through the construct.
Among materials studied GelMA is at the forefront.[] Early
studies using GelMA as a bioresin highlighted its use in cre-
ating hemisphere constructs[] and observing cell response
and interaction with local geometry.[] Although one of the
biggest challenges in using naturally based soft biomaterials
for vat polymerization, is the fabrication of perfusable net-
works, researchers have begun using vat polymerization in an
eort to introduce vascular networks within bio fabricated con-
structs. Recently, Wadnap etal.[] investigated using GelMA in
a projection stereolithography to fabricate D Y-shaped tubular
constructs with living cells encapsulated. However the diam-
eter of the construct was  mm, which is relatively low for
lithography-based applications.[] Later the group developed a
D four-branch vascular-like construct (Figure9c) with cell via-
bility of % after  h incubation.[] Interestingly, He etal.[]
used DLP-based printing with GelMA to print an “all-in-one”
gel microsphere system for adipose-derived stem cell (ADSC)
culture, with microspheres ranging in diameter from  to
 µm (Figure a). The microspheres enhanced cell expan-
sion, convenient and harmless passage, quick and complete
harvest, ecient cryopreservation and transportation. The
microspheres were also capable of creating functional mac-
rotissue. In an eort to aid the structural integrity of GelMA
during the fabrication process and create lager constructs
CaCO may be incorporated into the resin, which is removed
after construct fabrication using HCl.[] This has allowed for
the formation of GelMA scaolds with aligned gelatin fibres
mimicking the collagen bundles. However using this method
cells are seeded after the fabrication process rather than being
encapsulated in the bioresin.[,] Recently, GelMA has been
employed by volumetric tomographic printing, whereby free-
floating structures were printed including functional hydrogel-
based ball-and-cage fluidic valves.[] Additionally, anatomically
shaped cell encapsulated constructs were fabricated and found
to support the synthesize of new-tissue matrix in vitro. Inter-
estingly, there was a need for postprinting processing when
GelMA was used as the bioresin for volumetric tomographic
printing, where the printed constructs were further immersed
into a secondary photoinitiator bath, followed by additional
photoirradiation. By contrast, a recent publication by Rizzo
et al.[] showed that when gelatin–norbornene (GelNB) was
used for volumetric tomographic printing, there was no need
for postprinting crosslinking. Although both GelMA and
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Adv. Mater. 2022, 34, 
Figure 8. Synthetic-based bioresins showing a) cell interactions with fractal topographies. i,iv) SEM images of the corresponding hydrogel structures
conjugated with fibronectin. ii,v) DAPI/actin stained fluorescent microscope images, and iii,vi) DAPI/actin stained fluorescent microscope images at
higher magnification, of murine myoblasts and human mesenchymal stem cells (scale bar = µm). Reproduced with permission.[] Copyright ,
ACS Publications, b) D patterns and architectures not achievable with extrusion-based printing systems, including (i–iv) woven mats, (ii–v) chain
mail designs, and (iii–iv) D lattice structures with arches oriented in the z-direction. Reproduced with permission.[] Copyright , IOPscience.
c) Entangled vascular networks displaying i) interpenetrating Hilbert curves, ii) bicontinuous cubic lattice, iii,iv) torus and (,) torus knot (scale bars:
mm), v) elaboration of a lung-mimetic design through generative growth of the airway, oset growth of opposing inlet and outlet vascular networks,
and population of branch tips with a distal lung subunit, and vi) the distal lung subunit is composed of a concave and convex airway ensheathed
in vasculature by D oset and anisotropic Voronoi tessellation. Reproduced with permission.[] Copyright , AAAS. d) Photograph of a printed
hydrogel containing the distal lung subunit during RBC perfusion while the air sac was ventilated with O (scale bar: mm). Reproduced with permis-
sion.[] Copyright , AAAS.

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GelNB uses gelatin as the parent polymer, the crosslinking
process is fundamentally dierent (chain-growth vs step-
growth polymerization), which might explain the dierence in
bioresin performance observed.[] Similar to GelMA, GelAGE
is a derivative of gelatin but instead of the addition of meth-
acryloyl groups, gelatine is reacted with allyl glycidyl ether
(AGE). GelAGE uses the thiol–ene chemistry to form carbon–
carbon double bonds. Bertlein etal.[] have researched the use
of GelAGE as a bioresin for vat polymerization using DTT as
crosslinker and Ru/SPS as the photoinitiator to obtain cell free
high shape fidelity constructs (Figureb).
Hyaluronic acid (HA), which is a naturally occurring ECM
component and FDA approved biomaterial[] has also been
used in vat polymerization.[] HA is widely used in biomedical
applications due to its many critical biological functions, such
as regulating cell adhesion, influencing cell proliferation and
dierentiation. Similar to GelMA a photocurable form of HA
has been developed though the addition of methacrylate groups
to the polymers backbone.
Glycidyl methacrylate modified hyaluronic acid (GMHA) has
been investigated for nerve tissue engineering and although
multilumen scaolds were printed it required the addition of
wt% of zinc acrylate to enhance the printability of the resin
and rigidity of the construct.[]
In recent years SF has shown to be a positive candidate as
a bioresin. SF is found in the silkworm Bombyx mori and is a
fibrous protein, which consists of a repeating pattern of amino
acids. Adding methacrylate groups to the side groups of SF
allows it to be photo-crosslinked and used in vat polymerization.
Kim etal. used glycidyl methacrylated (GMA) SF to construct
complex organ structures, including the trachea, heart, lungs,
and brain with excellent structural stability and reliable bio-
compatibility (Figure10b).[] Additionally, SF-GMA constructs
built using DLP has shown promise in vivo with regards par-
tially defected trachea rabbit model. New cartilage like tissue
and epithelium was found surrounding the transplanted
SF-GMA hydrogel.[] Recently, Ajiteru et al.[] developed a
printable bioresin for neural tissue engineering, through cova-
lent reduction of graphene oxide (GO) by glycidyl methacrylated
silk fibroin. The resin was capable of printing complex shapes
as shown in Figure a. The printed hydrogels demonstrated
superior mechanical, electroconductive, and neurogenic prop-
erties than silk fibroin and supported Neuroa cell proliferation
and viability. More recently the same group developed a mag-
netic hydrogel based on SF-GMA and iron oxide for use in a
magnetic bioreactor system for the dierentiation of myoblast
cells.[]
An overview of natural bioresin formulations in literature
are given in Table3. As evident from Tables and, using vat
polymerization synthetic-based bioresins are at the forefront of
current research for tissue engineering and biomedical appli-
cations. This distinct lack of natural-based bioresins is due to
the complex nature of the printing process and the demanding
material features required to withstand the printing process
as discussed throughout this paper. Therefore, to advance this
technology further in the field of regenerative medicine there is
a pressing need to develop novel natural-based bioresins, which
are capable of being printed using vat polymerization, and
Adv. Mater. 2022, 34, 
Figure 9. Natural-based bioresins showing a) printing of microspheres (MSs). i) Schematic diagram of layer-by-layer printing of GelMA MSs, ii) fluo-
rescence images of GelMA MSs ranging in diameter, and iii) diameter analysis of GelMA MSs. Reproduced with open access.[] Copyright , Wiley,
b) i,ii) One-layered  wt% GelAGE constructs. Reproduced with open access.[] Copyright , Wiley, c) i) Fabricated of D vascular-like construct
using GelMA, ii) cell viability, and iii) the hollow cross-section. Reproduced with permission.[] Copyright , Elsevier.

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support cell encapsulation and matrix formation over relevant
culture periods.
7.3. Multimaterial and Gradient Structures
One of the major drawbacks of vat polymerization compared
to other biofabrication methods such as extrusion-based bio-
fabrication is its use of a single material. For example with
extrusion-based printing several print heads can be installed
into the printer, each loaded with a dierent biomaterial ink,
to allow for printing a continuous multimaterial layered struc-
ture.[,,] However, with vat polymerization printing mul-
timaterials is more challenging as the vat can only contain one
material. To overcome this and allow for a diverse range of
materials within a single structure one of two methods may be
employed, either i) perform a complete vat exchange, whereby
the vat is removed and changed with a vat containing a new
material or ii) the existing material is removed from the vat,
the vat is cleaned and refilled with new material. However, both
of these techniques is labor intensive and add increased print
time to the construct. Furthermore, as the vat can only con-
tain one material at any given time, this means each layer can
only be made up of a single material and it is not feasible to
create multimaterial layers. The main challenge to be addressed
in creating a multimaterial construct is to reduce the material
waste and increase cleaning eciency during the resin tank
switching process. In a top-down approaches the vat changing
system can waste some of the resin and due to its nature of
the entire construct being submerged in the resin the proba-
bility of contamination can be significant.[,] Zhou etal.[]
investigated the bottom-up projection approach whereby vats
containing dierent materials were automatically changed to
alternate resins. By using this approach, the amount of mate-
rial used was minimized as only small amounts of material
were require compared to top-down. Additionally, contamina-
tion between resins was also lowered as only a small surface
area of the construct was submerged. Ge etal.[] employed a
similar concept to create high-resolution constructs with multi-
material architectures. Using this multimaterial technique, the
group printed D shape memory polymer grippers as shown in
Figure11b. By placing the multiple materials in dierent posi-
tions, could alter the mechanisms of the grippers in response
to a stimuli, to enable dierent functionalities, including
closing the grippers for grabbing objects and opening gripper
for releasing objects. More recently, Kim et al. D printed a
bilayered silk fibroin bioresin for tissue engineering using
DLP.[] The shape of the construct could undergo a revers-
ible shape transformation upon transfer between pure and salt
water. Additionally, two cell types (chondrocytes and turbinate-
derived mesenchymal stem cells) were incorporated into the
construct to mimic the heterogeneity of the trachea, with cell-
laden constructs then implanted into damaged rabbit trachea.
The D printed constructs were found to integrate with the host
trachea, and both epithelium and cartilage were formed at the
predicted sites. Using the combination of DLP and D printing
the authors were able to print multilayered cell cylindrical con-
structs, which could not be solely facilitated by DLP.
In these multimaterial approaches it is essential to con-
sider the force being applied to the construct when cleaning
between each vat change, in particular soft bioresins for tissue
engineering applications and cell laden resins. Suri etal.[]
demonstrate “proof-of-concept” to visualize the feasibility
of gradient formation using DLP. A gradient was prepared
from the same bioresin incorporating red and green fluo-
rescent microparticles in opposite directions to demonstrate
Adv. Mater. 2022, 34, 
Figure 10. Silk-based bioresins showing a) i) schematic of bioresin for neural tissue engineering, through covalent reduction of graphene oxide (GO)
by glycidyl methacrylated silk fibroin and printed. ii) Alveolar structure and image of a transverse section of a human spinal cord. The impression
is indicated by gray matter and dorsal roots using these materials. Reproduced with permission.[] Copyright , ACS Publications, b) i–iv) Cell
distribution inside the Sil-MA hydrogels. The Sil-MA hydrogels with i) a design of the letter HL (the logo of Hallym University), ii) a shape of human
brain were printed out with PKH-labeled cells only, iii) the Sil-MA hydrogels with the letter HL, and iv) a shape of winding trachea were printed out
with PKH-labeled cells (green) and PKH-labeled cells (red); (from left to right) CAD images, printed images, fluorescence images by (i) confocal
or (ii–iv) single plane illumination microscopy (SPIM) microscope, and merged images of fluorescence and CAD images. Scale bar indicates mmon
CAD images and  mmon printed images. Reproduced with open access.[] Copyright , Nature Publishing Group.

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Table 3. Natural-based bioresin compositions for tissue engineering and biomedical application.
Material Printer set up Wavelength Initiator/crosslinker Photoabsorbing dye/
inhibitor
Coating
material/
additive
Cells Time point Functional outcomes Feature size Ref.
Cell free
GelAGE EnvisionTec,
Perfactory
Ru (. × −)/
SPS (× −),
DTT (allyl:SH=:)
Strand:
 µm
[]
SF/Gelatin DLP  to 
nm
Riboflavin (×
−)
Strand:
 µm
[]
Cell seeded
GelMA/CaCO PSL UV light Irgacure 
( w/v%)
HMBS (.% w/v),
TEMPO (.%
w/v)
Human avascular
zone meniscus cells
Week  (Ex
vivo)
Targeted for meniscus regeneration, showed
formation of fibrocartilage-like tissue and
integration with tissue in an explant model.
Micropatterned scaolds are non-cytotoxic and
facilitated organized cellular alignment.
Strand:
 µm
Pore:
 µm
[]
GelMA/calcium
carbonate
microparticles
DLP UV light Irgacure 
( w/v%)
HMBS (. w/v%),
TEMPO (.
w/v%)
Immortalized human
umbilical vein
endothelial cells
Day  Interconnected pores enabled uniform
cell seeding distribution. The seeded
cells were able to proliferate to reach
confluency, and maintained their
endothelial phenotype.
Pore:
µm
[]
GMHA UV light Irgacure 
( w/v%)
TEMPO
(. w/v%)
wt% of
zinc acry-
late, laminin
coating
Schwann cells Day  Focused on shape and architecture, as well as
two dierent gradients within a scaold. Cell
functionality is limited to only cell adhesion after
short culture period.
Pore:
≈ µm
[]
GelMod PP  nm Irgacure  Human adipose-
derived stem cells
(ADSCs)
Day  Evaluated adhesion, proliferation, and adipo-
genic dierentiation of ADSCs. Formation of
fat vacuoles and cell buoyancy evaluated
as measure of adipogenesis.
Pore:
≈ µm
[]
GelMA DLP  nm LAP (.% w/v) Phenol red
(.% w/v)
hADSC Day  Formation of gel microspheres with rough
surface architecture and tailorable mechanical
properties as platform for cell expansion.
ADSCs expanded using the microspheres were
able to dierentiate into both osteogenic and
chondrogenic lineages.
Diameter:
 µm
[]
Cell encapsulated
GelMA Smartphone-
enabled DLP
Ru ( × − )/
SPS( × − )
CC Day Justshowedviability over  days culture time
point and change in cell morphology.
Functionality assessment not performed.
[]
GelMA DLP  nm Irgacure 
(. w/v%)
NIH T mouse
fibroblasts
Day  No functionality assessment, just showed
viability ≈% post printing.
Wall:
 µm
Diameter:
mm
[]

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Material Printer set up Wavelength Initiator/crosslinker Photoabsorbing dye/
inhibitor
Coating
material/
additive
Cells Time point Functional outcomes Feature size Ref.
GelMA DOPsL  nm Irgacure 
(.– w/v%)
Fibroblast Day  No functionality assessment, cell viability
evaluated to be % after  days post-printing.
Wall:
µm
(designed)
[]
GelMA DOPsL nm Irgacure 
(. w/v%)
NIH-T murine
embryonic fibroblasts
(Ts) and
CH/ T/ murine
mesenchymal pro-
genitor cells (T/s)
Day  Evaluated the eect of construct geometry and
pattern on cell proliferation and morphology.
Cell viability >%observed post printing, but
subsequent functionality not evaluated
Feature:
≈ µm
[]
GelMA Volumetric
tomography
 nm LAP (.%
w/v), post photo-
crosslinking Ru
(. × − )/
SPS(× − )
Matrigel Equine-derived ACPCs,
bone marrow-derived
MSCs, endothelial
colony forming cells
(Ecco)
Day 
(Bone)
Day 
(Meniscus)
Functionality of ECFC assessed by formation
of vessel length, where volumetric bioprinting
resulted in higher vessel length and interconnec-
tivity compared to control D culture condition.
Functionality of ACPCs evaluated via matrix/
tissue formation—glycosaminoglycans, collagen
Type I & II. Encapsulated ACPCs were able to
secrete matrix after  days in culture which
resulted in increased overall
compressive strength.
 µm[]
Gel-NB PP  nm DAS ( × −) and
DTT (allyl:SH= :)
L mouse fibroblast
cells
Week  Evaluated morphology and migration
of Ls embedded within printed constructs.
Cells demonstrate rounded morphology in stier
regions, whereas are more extended and
migratory in softer areas.
Pore: 
µm
[]
Gel-NB Volumetric
tomography
 nm LAP (.% w/v)
DODT, PEGSH,
PEGSH
CC, human
neonatal fibroblasts
(NHDFs)
Day  High cell viability >%observed after
 week of culture. Cell spreading and prolifera-
tion visualized using bright-field microscopy.
 µm[]
Silk graphene
oxide
DLP UV light LAP mouse neuroblastoma
cell line (Neuroa)
Day  Neuroa cells were able to proliferate within the
printed constructs, as well as being
neurogenically active. Functionality assessed via
immunofluorescence image analysis, as well as
RT-qPCR for neuronal markers
Feature:
>mm
[]
GelMA DLP  nm Ru/SPS Tartrazine Human mesenchymal
stem cells (hMSCs)
Day  hMSCs able to undergo chondrogenic,
hypertrophic and osteogenic dierentiation
after  days in culture Histological assess-
ment showed key cartilage and bone markers
(collagen type II and mineral deposition) for
all chondrogenic, hypertrophic and osteogenic
dierentiation protocols.
 µm[]
Table 3. Continued.

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Material Printer set up Wavelength Initiator/crosslinker Photoabsorbing dye/
inhibitor
Coating
material/
additive
Cells Time point Functional outcomes Feature size Ref.
Silk-GMA DLP  nm LAP (.%) NIH/T mouse
fibroblast cells or
chondrocyte (In vitro)
Chondrocytes (In vivo)
Week  (In
vitro), Week
 (in vivo)
Long term functionality of chondrocytes evalu-
ated via glycosaminoglycan production and
expression of chondrocytic specific gene–Col-
lagen Type II, Sox, Aggrecan and Collagen Type
X. In vivo subcutaneous implantation study of
chondrocytes-laden printed construct showed
histological characteristics of cartilage tissue
with proteoglycan and collagen expression.
The chondrocytes-laden printed constructs
were further transplanted to a rabbit trachea
and showed good growth and integration with
the surrounding tissue, as well as neocartilage
formation.
Feature:
>mm
[]
Silk-GMA DLP  nm LAP (.% w/v) Human chondrocytes
and turbinate-derived
mesenchymal stem
cells
Week  (in
vitro), Week
 (in vivo)
Using D-bioprinting to control shape of printed
construct and also able to spatially localized
two dierent kind of cells. Bioprinted trachea
mimetic constructs were transplanted into a
rabbit trachea. Light stenosis observed, bio-
printed trachea able to integrate with the periph-
eral trachea and formation of new epithelium
that is similar to normal trachea obtained.
[]
Silk-GMA DLP  nm LAP (.% w/v) NIH/T mouse
fibroblast cells and
chondrocytes
Week  Long term in vitro functionality of chondrocytes-
laden printed constructs demonstrated by cell
organization and extracellular matrix production
(proteoglycan and collagen).
Feature:
 µm
[]
GelAGE: allylated gelatin, SR: silk fibroin; GelMA: gelatin methacryloyl; GMHA: methacrylate-modified hyaluronic acid; GelMod: methacrylamide-modified gelatin; Gel-NB: gelatin-norbornene; Silk-GMA: glycidyl meth-
acrylated silk fibroin.
PSL: projection stereolithography; DOPsL: dynamic optical projection stereolithography; DTT: dithiothreitol.
Table 3. Continued.

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gradient feasibility. However, it must be noted that as the same
material was used throughout the construct which eliminates
many challenges such as stitching between layers of dierent
materials. In another study spatial D layer-by-layer was dem-
onstrated by patterning distinct layers. In this study NIH/T
cells were encapsulated into a PEGDA bioresin however again
each layer consisted of the same material with cells stained
with either CellTracker CMFDA (green) dye or CMTMR
(orange) dye.[] Although this demonstrates the possibility
of spatial dierences further investigations are required for
Adv. Mater. 2022, 34, 
Figure 11. Use of multimaterials in vat polymerization. a) i) Schematic diagram showing vat polymerization setup of multimaterial dECM-based liver
cancer tissue platform and ii) merged fluorescence and paired bright field images showing the tracked HepG in dierent conditions. Reproduced
with permission.[] Copyright , Elsevier, b) i) Schematic of process of fabricating a multimaterial structure based on PµSL, ii) demonstrating D
printing, transition between as printed shape and temporary shape of multimaterial grippers. Reproduced undetr the terms of the CC-BY . license.[]
Copyright , The Authors. Published by Nature Publishing Group c) Set of photopatterned designs used to form the Mona Lisa using i) AF-
ovalbumin-PEG-thiol (green), ii) AF-ovalbumin-PEG-thiol (yellow), and iii) AF-ovalbumin-PEG-thiol (red). Scale bar = µm. Reproduced with
permission.[] Copyright , Elsevier. d) D printed developing cardiac tissue mimic composed of iPSC-CMs showing CellTracker Green-stained
HUVEC:T/ (:) cells growing from the sinus venosus to the capillary bed region in response to photopatterned + QK peptide over  h. Red
outline denotes printed region of the sinus venosus. Scale bar = µm. Reproduced with permission.[] Copyright , Elsevier.

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2107759 (30 of 35) ©  The Authors. Advanced Materials published by Wiley-VCH GmbH
multimaterial constructs using the layer-by-layer approach.
When using natural-based multimaterials in vat polymeriza-
tion rather than using a layer-by-layer approach researchers
often simply use a projection technique. Using this projec-
tion technique, the height of the construct is the thickness
of the cure depth, i.e., only one layer. This process has been
used to develop hepatic model, which was achieved by first
illuminating the resin with the desired pattern, removing the
uncured material, and then followed by pipetting a second
resin in the empty space of the first layer. By changing the
bioresin in this manner a functional hepatic tissue model pri-
marily based on GelMA, consisting of human-induced pluri-
potent stem cells-derived hepatic progenitor cells (hiPSCs-
HPCs) in one region and human umbilical vein endothelial
cells (HUVECs) and the ADSCs in another complementary
region was printed.[] The same group showed by using
GelMA and GelMA/HA-based bioresins, multiple cell types
mimicking the native vascular cell composition were encap-
sulated directly into hydrogels with precisely controlled dis-
tribution, which resulted in endothelial cells forming lumen-
like structures spontaneously in vitro.[] Recently, the group
used the same materials and patterning technology to print
sequential cell encapsulated materials to develop a cancer
tissue model,[] striated heart and lobular liver structures.[]
The cancer model was printed with regional stiness and nod-
ules mimicking the stiness of a cirrhotic liver. This regional
stiness enabled the visualization of HepG stromal inva-
sion from the nodule with cirrhotic stiness[] (Figurea).
More recently projection-based printing has shown capability
of generating well defined multiregion constructs by using
individual fluorescently labeled ovalbumin-PEG-thiols and
successively photopatterning three distinct digital patterns,[]
as shown in Figurec. Using this technique allowed for the
precise control over spatial distribution and concentration
of bioactive moieties, and was used to show the ecacy of
photo-conjugated QKpeptide to direct endothelial cell migra-
tion within a cellularized tissue construct[] (Figured). This
recent work demonstrates the significant advancements and
opportunities in the use of vat polymerization as a biofabri-
cation technique for biomedical applications such as complex
architecture or in vitro disease modeling.
8. Future Directions and Conclusions
The application of vat polymerization for the biofabrication of
tissue scaolds presents a significant opportunity and its appli-
cation is continuously growing. Although vat polymerization
is a well-established technique, its use as a biofabrication
method in the field of tissue engineering has been hindered by
the lack of appropriate bioresins that can withstand the printing
process, while allowing cell encapsulation and proliferation.
The layer-by-layer approach of DLP and SLA requests robust
materials to achieve a successful print, however for tissue engi-
neering applications low modulus materials are required for
nutrient diusion and cellular proliferation. Particularly as can
be seen from literature, many studies employing vat polymeri-
zation are currently lacking the encapsulation of cells directly
into the scaold, rather many studies seed cells on the surface
of the scaold.[] Improvements in the technology and material
development will help to overcome the current limitations that
prevent the biofabrication of tissue scaolds. Combining the
knowledge of chemistry, materials science, and biology is key to
developing materials for tissue engineering applications. Recent
studies have also established quantized models as frameworks
to predict and then analyze printability of materials in relation
to the targeted applications.[] Therefore, as outlined in this
review, understanding the key influences of the bioresin compo-
sition, along with the important requirements of vat polymeriza-
tion technique equips researchers with critical knowledge to aid
the development of novel bioresin formulations.
The recent development of volumetric tomographic printing
holds distinct advances in vat polymerization as it overcomes
challenges in traditional layer-by-layer techniques such as poor
surface quality and slow fabrication speeds. A key advantage of
this platform is the ability to print high viscosity fluids, which
are typically challenging with layer-by-layer printing setups.
Moreover, this technique provides a novel opportunity for fabri-
cating low modulus hydrogel structures for tissue engineering
applications, as the printed object remains suspended and sta-
tionary during the print, thus minimal forces are exerted on the
printed object. Utilising this fabrication techniques for tissue
engineering application allows for the construct to be printed in
a time frame ranging from seconds to tens of seconds, which
is far superior to DLP/SLA, opening new avenues for upscaling
the production of hydrogel-based constructs.[] These advan-
tage holds novel opportunities to pioneer bioresins for use in
high speed, organ-scaled tissue engineering applications.
A current major limitation with vat polymerization, is its
restriction to the use of one resin at any given time and thus
creating multimaterial constructs represents a significant chal-
lenge. Therefore, an opportunity lies in developing an auto-
mated system to change resin vats or remove uncured material
and substitute with another. This will allow for the build of more
complex multicellular scaolds, which better mimic the natural
environment in vivo. Vat polymerization, particularly volumetric
tomography printing, provides control over the architecture
and structure of scaolds that cannot be achieved with tradi-
tional biofabrication techniques such as extrusion. However, to
date, relatively simple architectures have been investigated for
tissue engineering approaches, with minimal naturally inspired
scaolds being investigated. Therefore, in the coming years,
through bioresin development, more complex architectures are
expected to be developed to mimic the architectures of the native
tissue. Furthermore, this presents researchers with a unique
opportunity to explore how scaold architecture can influence
cellular adhesion and migration. With future developments in
bioresins and printer technology, may allow for the production
of organ-on-a-chip devices that combine environment and archi-
tecture to create realistic human models. It is envisioned that
these models could greatly accelerate drug discovery with the
eventual elimination of animal models.
Acknowledgements
The authors acknowledge funding from the New Zealand Health
Research Council (Sir Charles Hercus Fellowship / (K.S.L.),
Emerging Researcher First Grant /, K.S.L.) together with the Royal
Adv. Mater. 2022, 34, 

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2107759 (31 of 35) ©  The Authors. Advanced Materials published by Wiley-VCH GmbH
Society Te Apārangi (Marsden Fast Start Grant MFP-UOO (K.S.L.),
Rutherford Discovery Fellowship RDF-UOO, T.B.F.W.), the Ministry
for Business, Innovation & Employment (UOOX; T.B.F.W.), and the
New Zealand Equine Trust Project Grant (T.B.F.W.).
Open access publishing facilitated by University of Otago, as part of
the Wiley - University of Otago agreement via the Council of Australian
University Librarians.
Conflict of Interest
The authors declare no conflict of interest.
Keywords
bioresin, biofabrication, bioprinting, regenerative medicine, vat
polymerization
Received: September , 
Revised: January , 
Published online: March , 
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Caroline A. Murphy is a Postdoctoral Fellow within the Christchurch Regenerative Medicine and
Tissue Engineering (CReaTE) group at the University of Otago, New Zealand. Before joining the
CReaTE group she completed her Ph.D. in Biomedical Engineering at the University of Limerick,
Ireland. Caroline’s expertise lies in the field of D-biofabrication technologies for tissue engi-
neering and regenerative medicine applications. Her primary research interests include expanding
the range of biomaterials for D bioprinting technologies, with a particular interest on the
development of scaolds with optimal biological and structural properties for cartilage tissue
engineering.
Adv. Mater. 2022, 34, 

www.advmat.dewww.advancedsciencenews.com
2107759 (35 of 35) ©  The Authors. Advanced Materials published by Wiley-VCH GmbH
Khoon S. Lim is an Associate Professor at the University of Otago, New Zealand. He leads the
Light Activated Biomaterials (LAB) Group and his research focuses on developing photopoly-
merizable hydrogel bioinks for D-bioprinting of functional tissues. He has received a number
of prestigious fellowships such as the Rutherford Discovery Fellowship from the Royal Society
of New Zealand, and the Sir Charles Hercus Health Research Fellow from the Health Research
Council of New Zealand. He currently serves as the Vice President of the Australasian Society for
Biomaterials and Tissue Engineering, as well as Council Member of the New Zealand Association
of Scientists.
Tim B. F. Woodfield is Professor of Regenerative Medicine, Department of Orthopaedic Surgery,
University of Otago, New Zealand. His research involves development of advanced bioinks,
spheroid bioassembly and additive manufacturing platforms applied to musculoskeletal regenera-
tive medicine. Tim is a Fellow, Biomaterials Science & Engineering (FBSE) and recipient of: Royal
Society of New Zealand Rutherford Discovery Fellowship; Australasian Society for Biomaterials &
Tissue Engineering (ASBTE) Research Excellence Award; Otago University Research Gold Medal.
Tim is President Elect, International Society for Biofabrication and former ASBTE President. He
sits on the TERMIS-AP Council and Editorial Boards of Biofabrication, APL Bioengineering and
Frontiers in Bioengineering & Biotechnology.
Adv. Mater. 2022, 34, 