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Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_980_Библиотеки_им_академика_М_И_Перельмана
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the material
Filament is led
Filament spool
The
and a pinch system to
and retract the filament
precise amount
wn
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Fig. 2 Schematic
illustration of Fused
to the extruder
Deposition Modeling
(FDM) with single
nozzle [11]
extruder uses torcpe
The heated filament is forced
out the heated nozzle at a
smaller diameter
feed
s.
A heater block melts the
filament to a useable
temperature.
The extruded material is laid do
on he model where it is needed.
The print head and/or bed is moved
to the correct X/Y/Z position for placing
The initial printing supply is often a thin lament, melted in nozzle and extruded
using pressure. Where thin material lament is employed as the printing supply, the
method can be referred to as fused lament fabrication (FFF). On the other hand,
the pellet or granules of material blends are utilized in precision extruding deposition methods [9]. These methods encompass material melting in a chamber and,
later on, extrusion via rotating screws (mini-extruder).
Multi-headed deposition systems are modied FDM setups with multiple nozzles for extrusion of different materials. Their development has enabled fabrication
of scaffolds consisted of layers with different compositions. Moreover, conjunct
FDM and bio-printing has enabled fabrication of cell-laden scaffolds, with cells
being encapsulated within the scaffold struts [10]. However, in terms of bio-printing
procedures, further investigations are warranted (Fig.2).
2.1 Prominent Features
Attainment of high porosity and a vast range of pore sizes (i.e., 250–1000μm) has
been enabled with FDM.Tailoring these factors contributes to reaching a balancing
ratio regarding the mechanical and biological properties of the scaffold. Additionally,
FDM-developed scaffolds often hold satisfactory rates of pore interconnectivity.
Porosity errors with FDM have been low, making it a relatively precise approach for
fabrication of porous constructs [12]. Moreover, upon proper material selection, apt
compressive strengths are within reach, and the mechanical properties of scaffolds
are not jeopardized during the procedure. With regard to the necessity of proper
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interlayer attachment, the template temperature must be maintained at degrees
below the solidication point of the working blend [13, 14]. Appropriate mechanical stability of FDM allows producing of scaffolds with higher porosities. Multiple
nozzles of certain FDM setup designs enable fabricating gradient scaffolds with
layer-specic material composition, pore shapes, pore sizes, and porosities. Of note,
addition of organic solvents is not obligatory for FDM, contributing to higher rates
of scaffold biocompatibility. Moreover, FDM is a known as a cost-effective additive
manufacturing method. On the other hand, high working temperatures are the chief
downside of FDM, conning the application of this technique when thermosensitive
biological agents are to be supplemented [15]. Moreover, the vertical and horizontal
pore sizes are often incongruent. When printing large particles, FDM bears chances
of poor surface nishing, compromised accuracy and a time-consuming procedure.
FDM is not ideal for fabrication of irregularly congured constructs [16]. With specic congurations, designing of temporary supporting structures may be required,
resulting in necessity of utilizing multi-nozzle setups. Since rheological features of
the working material dictates the printing resolution, a restricted spectrum of polymers can be utilized for FDM [17]. Such materials are required to offer a viscosity,
low enough that does not interfere with the extrusion procedure and high enough to
preserve the scaffold structure [18]. In many settings, supplementary rheological
modiers, while remaining aware of their impact upon scaffold biocompatibility,
can be implemented [18]. In order to obtain an optimal printing resolution, the rheological features of the working blend and the nozzle diameter can be tuned. Given
this, nozzle clogging must be accounted for while making procedural adjustments.
H. S. H. Boroojeni et al.
2.2 Employed Material
Thermoplastic materials are often applied in FDM. PCL is the most commonly
employed polymer for FDM-based scaffolds, owing to its low melting point of
60°C.Poly(lactic-co-glycolic acid) (PLGA) is another highly utilized material with
FDM.However, its high melting point can perplex the procedure. PLA, ethylene
vinyl acetate (EVA), poly(methyl methacrylate) (PMMA), and ABS are the other
material options for FDM.Polymers with high melting points, such as polyether
ether ketone (PEEK), have also been used. However, shrinkage, warpage, and
delamination phenomena, consequent to sharp temperature variations upon extrusion, can play detrimental roles [19]. FDM has been widely used to developed
polymer- ceramic composite constructs. Of note, FDM is an inappropriate technique
for fabrication of scaffolds with high mineral contents, due to higher chances of
nozzle clogging as the blend viscosity increases. Therefore, nozzle diameter and
size of employed particles must be prudently correlated. Owing to incongruent
shrinkage of ceramic and polymeric contents of composite blends, secondary pore
formation can be expected. In such cases, it is crucial to be aware of plausible alternations in terms of mechanical properties and perform compensatory approaches, if
required [19].
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2.3 Clinical Application
FDM has been applied in bone contouring and bone regeneration. Goh etal. [20]
applied FDM-fabricated PCL scaffolds, aiming for socket preservation. Probst etal.
[21] fabricated PCL-TCP scaffolds as patient-specic cranial degradable implants
with FDM.Their 6-month follow-up data indicated feasible rates of bone formation
and implant integration. Han etal. [22] applied 3D-printed PCL scaffolds to recontour maxillary defects. Their implanted scaffold degraded over a 3years’ period of
time, whereas the amount of bone formation was mentioned to be unacceptably low.
Patient-specic PLA auricular reconstruction models have also been developed
using FDM [23]. FDM has been most commonly employed for development of
prototyping models that allow preoperative bending of titanium plates, surgical
training, and defect analysis [24]. The notably compromised accuracy of FDM
method compared to other manufacturing methods is highly critical in preoperative
decision-making and may lead to sole educating applications in many sensitive
cases [25]. However, even in terms of educational applications, FDM-fabricated
models lack optimal haptic feedback [25].
Patient-specic drugs releasing mouth guards have also been fabricated using
FDM [26]. Using melt extrusion printing, the bone implant can be fabricated made
up of printable materials, such as ABS and PLA [27–30]. Then, the molding and
casting procedure is performed to produce implants made up of PMMA.Although
these implants are fabricated through indirect methods, proper accuracy of the
duplication procedure has been reported [31]. Application of FDM-fabricated PEEK
prothesis has also been shown [32].
2.4 Low-Temperature Deposition Manufacturing (LDM)
Low-temperature deposition manufacturing is another extrusion-based method,
with principles similar to FDM.It was rst developed in the Tsinghua University. In
LDM, the scaffold is fabricated within a chamber at temperatures below 0°C, and
the nal product is freeze-dried for removal of the frozen solvent. In addition, phase
serration in LDM elevates the porosity of scaffold [33]. Given that LDM is based on
low-temperature settings, biological molecules are less likely to be impacted.
Synthetic polymers, natural polymers, and composite materials have been successfully utilized in LDM for scaffold fabrication [34]. Moreover, gradient structural
properties have been obtained with LDM, as well [35].
Higher solvent concentrations have been indicated to deate nal pore sizes.
Considering the role of freezing in interlayer adhesion and fusion in LDM chambers, extrusion of already frozen struts, as a result of critically low nozzle temperatures, can lead to layer detachment. On the other hand, higher nozzle speeds and
diminished extrusion rates can result in reduced diameter of printed struts and thus
lead to increased chances of strut breakage [36].
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H. S. H. Boroojeni et al.
2.5 Powder Melt Extrusion
FFF and pellet melt extrusion techniques are restricted in terms of utilizing new
material blends. This is due to the necessity to form the material as pellets and bers
initially and, later on, implement them into manufacturing systems. In powder melt
extrusion (PME), also referred to as one-step direct powder printing, the initial
material powder is employed. However, PME technology offers compromised accuracy, resolution, layer consistency, and printing quality compared to FFF [37]. PME
in majorly benecial due to being deprived of preprocessing phases for preparation
of feeding stuck, used in FFF and pellet printing [31]. PME can be an appropriate
option when it comes to drug delivery [38]. Poly-lactic acid (PLA), High Impact
Polystyrene (HIPS), and acrylonitrile-butadiene-styrene (ABS) scaffolds have been
fabricated via PME.
3 Computer-Aided Wet-Spinning
Aiming for porosity requisites in scaffolds preparation, different approaches, including AM-based hybrid systems, have been investigated. They include incorporation
of pore-forming methods, freeze-drying, or wet-spinning (WS) [39].
Computer-aided wet-spinning (CAWS) is a recently emerged method wherein
constructs are fabricated as a result of polymeric solutions being jetted into laments and rested upon and solidied as laments on a deposition substrate placed
inside of a coagulation bath compartment. This method has allowed manufacturing
of local and/or scattered porous inner structures [39]. Adaptation of AM principles
by WS has contributed to promoting manufacturing automation and better adjustment of scaffolds’ architecture on macro- and microscales. Therefore, improved
reproducibility and more diverse scaffold designs can be attained. In contrast to
polymeric scaffolds fabricated via other melt extrusion-based AM methods, CAWS
bers are spongy and of hybrid porosity. Their spongy morphology has been
ascribed to the phase inversion process that impacts polymer solidication and provides the polymeric ber matrices with micropores. Moreover, nal products are
hybrid polymeric structures that are both globally macroporous and locally microporous. Other melt extrusion-based AM methods, such as FDM, can produce relatively dense nonporous polymeric matrices that do not resemble CAWS-manufactured
constructs [40]. CAWS has not only borrowed the accurate adjustability of external
morphology, macroporosity, and exible material selection of AM techniques, it has
also endowed WS advantages, endowing the polymeric matrices with microporosity
and drug-releasing capabilities [40, 41].
3.1 Prominent Features andEmployed Materials
Biomedical applications of wet-spun polymeric structures, such as sustained and
controllable drug-releasing systems, extracorporeal blood treatment, and tissue
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engineering, have been exhaustively investigated [40]. Initial biomedical applications of WS were rooted in their feasibility in terms of producing bers, composed
of natural polymers, that were prone to thermal degradation with other spinning
techniques [42, 43]. Several studies have reported fabrication of tissue engineering
scaffolds with indiscriminately oriented wet-spun bers. These scaffolds were
observed to have a highly interconnected and microporous architecture, favoring
cell adhesion, proliferation, and spreading [44–48]. Of note, wet-spun bers possess
versatile drug-loading methods and adjustable releasing kinetics. Thus, they can be
one of the major contributors when designing tissue engineering constructs with
drug-releasing purposes [43, 49]. CAWS is chiey advantageous in terms of developing functionalized well-structured brous constructs. This is attained through
using supplementary bioactive agents to the primary polymeric solution and provides the manufacturer with drug-loaded scaffolds, fabricated through single-stage
additive manufacturing that obviates the need for post-fabrication treatments,
implementation of more complex technologies, etc. [43]. Since a variety of therapeutics, such as antibiotics and chemotherapeutics, can be easily incorporated into
wet-spun bers, their possible choice of composition has broadened into synthetic
sources, as well [49]. Wet-spun biodegradable polymeric scaffolds can act as carriers for bioactive agents (e.g., chitin [50], chitosan [51, 52], polyacrylonitrile [53],
poly(, -lactic acid) [54], poly(-lactic acid) (PLLA) [55, 56], PLGA [57], or
*PCL [44, 45]). In addition, attempts have been made to assemble wet-spun bers
into degradable macroporous scaffolds [43]. A typical WS setup includes a viscosetype spinneret, a dope, a coagulation bath, and rollers with lament winding, drawing, and drying purposes [58]. In a CAWS unit, the working polymeric solution is
placed within a programmable syringe pump that ejects the material from a needle
tip at a constant feed rate, into a coagulation bath, in a layering fashion [41]. The
spatial motion of tip is programed by the CAD signals, and depositions of each layer
follow a predened lay-down trajectory within the coagulation bath, until the nal
product is obtained [40].
CAWS has been demonstrated as an apt method for fabrication of polymeric
constructs with predened geometry and distribution of pores on a micrometeric
scale, as well as customized macroscale dimensions. Layered scaffold prototypes
composed of biodegradable synthetic polyesters (e.g., PCL and poly(methyl methacrylate) [41, 59]), biodegradable natural polyesters (e.g., poly[(R)-3hydroxybutyrate- co-(R)-3-hydroxyhexanoate)] (PHBHHx)), and polyesters with
combined natural and synthetic sources (e.g., chitosan/poly(g-glutamic acid),
PHBHHx/PCL blends [39, 60, 61]) have been made with CAWS. In terms of
CAWS-fabricated drug-loaded scaffolds, Puppi etal. developed the AM star poly(ecaprolactone) (*PCL) scaffolds loaded with levooxacin. Morphological, thermal,
and mechanical characterization of their scaffolds indicated that the neither manufacturing procedure nor the nal properties of constructs were negatively impacted
by drug-loading. Their *PCL scaffolds were also capable of sustained antibiotic
release over a period of 5weeks invitro [39].
Fiber porosity can be adjusted within a nano-/microscale spectrum through tuning of CAWS pertaining technical parameters in phase inversion and/or deposition
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stages. It is backed by accumulating evidence and theory of non-solvent-induced
phase separation that polymer concentration, solvent/non-solvent systems, solution
feeding rate, and deposition velocity can be tailored in ways to provide predened
porosities that can impact other scaffold properties, such as biodegradation rate,
drug release kinetics, cell-scaffold interactions, and lavage management [62, 63].
Given its adjustability of parameters, CAWS method has demonstrated a great
potential for nanotopographic surface manipulations. Surface tailoring enables
enhancing of protein-mediated cell adhesion and behaviors [63, 64]. However, it
must be noted that with increased porosity, mechanical properties of scaffolds are
prone to becoming compromised. The spongy structure of CAWS bers that form
each layer results in overall compromised mechanical properties compared to scaffolds of similar porous structures with dense struts, fabricated by other melt extrusionbased AM methods [40, 65]. Puppi etal. fabricated poly(ε-caprolactone) (PCL) and
PCL/hydroxyapatite (HA) scaffolds using CAWS. Their scaffolds were easy to
reproduce, in terms of internal architecture, and highly porous and had well- oriented
bers of 200–250 μm diameters. Reports of mechanical characterization of their
scaffolds suggested that the architecture and HA loading impacted compressive
modulus and construct strength. Feasible cellular behavior and bone mineralization
were reported [41]. Of note, anatomical *PCL scaffolds fabricated via CAWS indicate a ne structural stability for bone regeneration in non-load-bearing rabbit radius
critical-sized defects [66]. The *PCL scaffolds did not require any further xation
upon implantation and were desirably integrated with the surrounding recipient beds.
A three-month follow-up of their scaffolds showed new bone formation [43, 67].
H. S. H. Boroojeni et al.
4 Inkjet Printing
The inkjet printing method is conducted in a noncontact reprographic manner.
Using inkjet printing, ink drops are employed to produce an image or character
based on the computer-uploaded digital data onto a substrate matrix [68–71].
Construct fabrication functions with a “bottom-up,” unlike the conventional “topdown,” approach. Inkjet printing can be also referred to as the “droplet-based printing” due the nature of solution or suspension droplets of nano-to pico-liter volume
range, being jetted from a nozzle [72, 73]. The method is rooted in upstream introduction of a small volume change that alters pressure and results in ejection of the
pending ink droplet [74, 75]. Inkjet printers are known for their high speed, functional accuracy, and their exible selection of available biomaterials. Their feasible
accuracy and resolution lie in adjustability of pertaining functional parameters, such
as droplet size and deposition rates [75, 76]. Inkjet printing can be conducted
promptly with low costs. For instance, the computer designs are directly printed
without the need for fabrication of a pre-mask component. Overall, the reduced
printing stages manifest inkjet procedure as a low-cost and relatively simple technique [77] and enable cell encapsulation within the printed material [74, 78]. In
addition, the noncontact nature of inkjet printing procedure lowers the risk of contamination [77, 79].
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4.1 Classifications
Two modes of ink deposition, namely, droplet formation mechanisms, are present
with inkjet technology that dissect this method into two major categories: (1) continuous inkjet (CIJ) and (2) drop-on-demand inkjet (DOD). The two techniques also
differ in terms of droplet size and distribution. DOD droplets have lower diameters
and are more evenly distributed [77].
4.1.1 Continuous Inkjet (CIJ)
Continuous inkjet (CIJ) printing refers to rapid jetting of a conductive uid ink that
forms a continuous stream of material [72]. Imposition of an electric eld onto the
ink electrically charges and deects the jet toward the collecting substrate, based on
CAD signals [72, 73, 80]. Given that the non-deected jets are retrieved to the
printer reservoir, contamination may occur despite CIJ’s rapid deposition rates [77].
CIJ printers can be categorized as binary deection, multiple deection, hertz, and
microdot inkjet printers, based on the method used for drop deection. Binary
deection systems print uncharged droplets simultaneously on collecting matrices
[81, 82]. On the other hand, multiple deection systems deect the charged droplets
onto the matrix and multiple dot positions per nozzle are attained [81, 83]. By virtue
of multiple deposition, multiple deection systems have higher printing speeds than
binary deection systems. Moreover, hertz CIJ systems are modications of the
binary deection systems with enhanced color printing capabilities, and the number
of deposited ink drops are dened by the volume of each pixel. Upon ejection of
differently sized droplets, only droplets of a diameter bellow a critical amount can
jet under the imposition of electrical eld. This manifests as a selective jetting process [81] that obviates the need for nozzle diameter reduction in newer systems [84].
4.1.2 Drop-on-Demand (DOD) Inkjet
Drop-on-demand (DOD) inkjet printers program the droplet ejection process [85].
Based on CAD, the jetting device of drop-on-demand inkjet printers actuates and
initiates droplet jetting. The ink deposition trajectory is accurately manifested via
CAD-based conduction of printhead or the substrate in X-Y-Z Cartesian coordinates
[72, 73, 80]. DOD inkjet printing can be chiey classied into thermal or piezoelec-
tric variants.
Thermal Inkjet
In thermal actuation, a heating element heats the printing head and vaporizes the
biomaterial, leading to droplet deposition [75]. This is followed by formation of
small bubbles that generate pressure pulses, resulting in droplet ejection in various
diameters, ranging from 10 to 150pl [74, 86–88]. The temperature gradient, current
pulse frequency, and working ink viscosity dictate droplet diameter [86, 88].
Piezoelectric Inkjet
The “piezoelectric inkjet printing” term reveals the presence of a piezoelectric
material, namely, specic solid materials that are mechanically deformed under
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electric eld imposition or the other way around, over the uid cavity of the system’s jetting device [89, 90]. Once CAD signals are received, the piezoelectric
material is programed to contract, leading to generation of voltage pulses that induce
pressure alternation and actuation of droplet ejection from nozzle [72, 74, 75, 91].
Polycrystalline piezoelectric ceramic placed in each nozzle allows generation of
transient pressure for ink ejection [92].
Electrostatic, Electrohydrodynamic, Solenoid Valve, andAcoustic Inkjet
Except for thermal and piezoelectric inkjet, other DOD inkjet systems are also available; however, their application has been limited [72, 73, 81, 86, 89]. For instance,
the electrostatic and electrohydrodynamic inkjet printing systems require working
with conductive inks. Moreover, solenoid valve inkjet printing bears the chances of
ejection of droplet sizes as large as 500μm or beyond that can compromise printing
resolution. In addition, acoustic printing requires an upside down substrate placement [72, 81, 93]. In electrically actuated inkjet systems, an electric eld is imposed
between the jetting device and the collecting substrate. The droplet is formed as a
result of complex interactions between surface tension ratio between ink and nozzle
and the applied electric eld. Upon signals being received by the printing head,
forces are balanced in favor of droplet ejection [85]. On the other hand, acoustically
actuated inkjet printers make acoustic radiation pressure, associated with ultrasound
elds that generate isolated droplets of identical diameters from an air-liquid interface. The droplet diameters and ejection pace can be adjusted by tuning of the ultrasound parameters (i.e., pulse, duration, and amplitude) [89, 94] (Fig.3).
Continuous
Binary Deflection
Multiple Deflection
Fig. 3 Inkjet Printers
Hertz
Microdot
Inkjet Printing
Drop-on-Demand
Electrostatic
Electrohydrodynamic
Solenoid-valve
Acoustic
Piezoelectric
Thermal
Substrate
Bioink
Heating
Element
Vapor
Bubble
Piezoelectric
Element
Nozzle
Droplet
Bioink
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4.2 Fields ofApplications
Previous to medical application of inkjet technology, it has had a widespread application in word processing areas, being an automation ofce tool. Moreover, it used
widely for printing of electronics in electronic and micro-engineering industries
[85, 95, 96]. More recently, the successful adaptation of the inkjet technology to
medical and biomedical engineering has brought its use into elds of drug screening, genomics, and biosensors [97, 98]. For instance, using inkjet printers, biological molecules, such as DNA, can be directed into high-density DNA microarrays,
despite their inherent fragility and without molecular degradation [99]. In addition,
active enzyme arrays for analyses have been generated using inkjet to deposit proteins substances, such as horseradish peroxidase, on cellulose paper [100].
4.3 Investigations andEmployed Materials
A exible selection of powder materials, including polymers, ceramics, proteins, and
cells, is available for inkjet printing [74, 101]. The rheological nature of the working
ink must be within a desirable spectrum. A viscosity of 5–20Pa.s is necessary for
continuous material ow. Viscosities below or beyond this range could cause high
ejection pressure or damp the necessary pressure wave, respectively [74, 77, 101].
In terms of bone regeneration, inkjet printing bioinks can impact not only the
invitro osteoblast differentiation but also invivo bone formation. Inzana JA etal.
printed 3D calcium phosphate scaffolds with collagen coating via thermal inkjet
printing. Upon 9 weeks implantation of scaffolds in critical-sized murine femoral
defects, implants were conrmed of osteoconductivity [102]. Three-dimensional
microgranular tricalcium phosphate scaffolds, fabricated by inkjet printing that
were chemically modied in buffer solutions, are also feasibly biocompatible [103].
Cooper GM et al. designed bone morphogenetic protein-2 (BMP-2) pathways
within microporous scaffolds via inkjet printing for implantation in murine calvarial
defects. The patterns of BMP-2 demonstrated cell differentiation invitro and tissue
formation invivo [104]. Another study, wherein collagen constructs were indirectly
printed via piezoelectric inkjet printing, suggested that high working temperatures
and constrained raw material input selection were alleviated. Their data demonstrated the role of intermediate-scale (i.e., >200mμ) pores and canals in ow and
mass transport and cellular scale (i.e., < 200mμ) surface texture of constructs in
cellular behavior, including surface attachment, proliferation, and migration. Since
iron-based alloys have high strength and corrode slowly, inkjet printing of them has
been studied. Inkjet 3D printing of mechanically milled Fe–30Mn (wt.%) powders
allows forming complex and customized products. Fe-based raw materials must be
machined into desired structures. This procedure has allowed direct processing of
such materials into scaffolds, with a reported open porosity of 36.3%. Electrochemical
corrosion tests suggested that the 3D printed Fe–Mn corroded faster than pure iron
constructs. Moreover, scaffolds demonstrated tensile mechanical properties analogous to native bone tissue, leading to decreased stress shielding risks. Fe–Mn
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scaffolds were reported to have more invitro cytocompatibility compared to plastic
tissue culture plates. Moreover, cell inltration into open pores was inspected [105].
Inkjet printing of mammalian cells has remained challenging, due to membrane
damage and cell lysis after sonication at 15–25kHz or during piezoelectric inkjet
printing [106]. On the other hand, although the nozzle tip temperature is locally
elevated up to 300°C for periods of microseconds in thermal inkjet printers [88],
the printed mammalian cells are documented to be exposed to a raised temperature
of 4–10°C above ambient temperature, for 2μs. This suits the 90% average cell
viability reported for thermal inkjet printing of mammalian cells [107]. Thermal
inkjet printers have also been documented to offer higher convenience regarding
modication, access, and maintenance than piezoelectric systems.
H. S. H. Boroojeni et al.
5 Stereolithography (SLA)
SLA was developed by Chuck Hull in 1984 and has been the rst dentistryimplemented additive manufacturing technique. It cures liquid material into 3D
scaffolds through exposure of a specied wavelength and intensity of ultraviolet
light (UV) onto primed photo-curable material. Once a thin solid layer is formed via
chemical polymerization, the printing platform descends to allow formation of the
second layer. This process continues in a layer-by-layer manner until the whole
structure is brought into form. SLA can manufacture polymer-ceramic composite
constructs and drug-loaded scaffolds [108, 109]. Through two irradiation trajectories, SLA can be conducted: the mask-based method and direct writing. The former
encompasses irradiation through a patterned mask for selectively exposing the predesigned patterns onto the photosensitive liquid polymer and triggers polymerization reactions in a spatially specic manner for generation of each layer. This
approach bears the cost of providing multiple masks. In order to avoid generating
multiple masks, liquids crystal display (LCD) or digital processing projection can
be used alternatively as a exible mask pattern. The other method, being the direct
writing SLA, uses concentrated UV beams to selectively solidify the polymer liquid
into the desired pattern. Irradiation may initiate from the above or the bottom of the
vat [110]. The mechanical properties and resolution of printed structures are highly
dependent upon the source of illumination that proceeds the process.
Most SLA systems use the single-photon polymerization process. On the other
hand, through two-photon polymerization, each present molecule engages with two
new photons into higher singlet states [111]. The two-photon polymerization
method allows submicron-scaled manufacturing of constructs and provide timewise benets [13] (Fig.4).
5.1 Prominent Features
SLA offers high-resolution structures, such that among all AM classications, SLA
has the most accuracy. Micro-SLA has also enabled provision of 20μm manufacturing accuracy [113].
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