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Chapter 40
Hepatic tissue engineeringAmanda X. Chen1,2,*, Arnav Chhabra2,3,4,*, Heather E. Fleming2,3 and Sangeeta N. Bhatia2,3,4,5,6,7
1Department of Biological Engineering, Massachusetts Institute of Technology, Cambridge, MA, United States, 2David H. Koch Institute for
Integrative Cancer Research, Massachusetts Institute of Technology, Cambridge, MA, United States, 3Harvard-MIT Health Sciences and Technology,
Massachusetts Institute of Technology, Cambridge, MA, United States, 4Institute for Medical Engineering and Science, Massachusetts Institute of
Technology, Cambridge, MA, United States, 5Wyss Institute for Biologically Inspired Engineering, Boston, MA, United States, 6Department of
Electrical Engineering and Computer Science, Massachusetts Institute of Technology, Cambridge, MA, United States, 7Howard Hughes Medical
Institute, Chevy Chase, MD, United States
Liver disease burden
Fatal liver disease accounts for B2 million deaths annu-
ally worldwide and has steadily increasing rates over the
years [1]. Liver failure can be divided into three major
categories: (1) acute liver failure (ALF) that presents as a
rapid loss of liver function in patients without preexisting
liver disease, (2) chronic liver disease due to metabolic
dysfunction, and (3) chronic liver failure accompanied by
tissue remodeling and scarring.
ALF is a rare syndrome with an annual incidence of
less than 10 cases per million people in the developed
world. In the United States, B2000 cases of ALF are
diagnosed each year [2]. It commonly develops in healthy
adults in their 30s. Patients with ALF usually present with
abnormal liver biochemistry, coagulopathy, and encepha-
lopathy. The causes vary geographically. Damage due to
drug exposure (e.g., acetaminophen) is the most common
cause in the West, while in large parts of the East, viruses
(e.g., Hepatitis A and E) are the most prominent cause of
ALF [3]. Clinically, ALF can be subdivided based on the
period of time between the appearance of jaundice and
onset of hepatic encephalopathy. Data from O’Grady
et al. proposed the following classification: hyperacute for
periods between 0 and 7 days, acute for periods between
7 and 28 days, and subacute for periods between 4 and 12
weeks [4]. In hyperacute cases the cause is usually acet-
aminophen toxicity or viral infection. Subacute cases that
evolve slowly often result from idiosyncratic drug-
induced liver injury (DILI). Even though patients with a
subacute presentation have less coagulopathy and enceph-
alopathy, paradoxically they have a consistently worse
medical outcome than those with a more rapid onset of
the disease [5].
Chronic liver disease develops on the background of a
constant injurious insult, either resulting from a metabolic
disorder or a number of etiologies that lead to widespread
tissue remodeling and pathologic deposition of extracellu-
lar matrix (ECM). Inborn liver-based errors of metabolism
are life-threatening conditions caused by genetic defects
in single enzymes or transporters and lead to blockade of
a specific metabolic pathway. While they can be accom-
panied by progressive fibrosis and cirrhosis, such as in
the case of α1-antitrypsin ZZ deficiency, hemochromato-
sis, Wilson’s disease, and hereditary tyrosinemia [6], the
liver parenchyma often remains intact. Some examples of
metabolic disorders with an intact parenchyma include
hypercholesterolemia, Crigler�Najjar syndrome, ornithine
transcarbamylase deficiency, organic acidurias, and
hyperoxaluria [7]. In a biopsy the lack of parenchymal
destruction often leads to a delayed diagnosis, exposing
the patient to sequelae. In all cases of liver disease the
lack of FDA-approved noninvasive biomarkers makes it
challenging to diagnose and treat liver diseases.
Chronic liver disease occurs in the setting of nonalco-
holic fatty liver disease (NAFLD). NAFLD is marked by
hepatic steatosis and is related to the presence of meta-
bolic syndrome in association with obesity, diabetes, and/
or arterial hypertension [8]. A subset of NAFLD patients
* These authors contributed equally.
737Principles of Tissue Engineering. DOI: https://doi.org/10.1016/B978-0-12-818422-6.00041-1
Copyright © 2020 Elsevier Inc. All rights reserved.
will develop signs of nonalcoholic steatohepatitis
(NASH), a more severe condition associated with lobular
inflammation and hepatocellular ballooning and can lead
to fibrosis and cirrhosis [9]. In NAFLD the liver is unable
to utilize carbohydrates and fatty acids properly, leading
to toxic overaccumulation of lipid species. These metabo-
lites induce cellular stress, injury, and death, which pre-
dispose the liver to sequelae such as cirrhosis and
hepatocellular carcinoma [10].
In the United States the number of NAFLD cases is
projected to expand from 83.1 million in 2015 (B26% of
the population) to 100.9 million by 2030 (B28% of the
population) [11]. An increasing percentage of these cases
are projected to be classified as NASH, rising from 20%
to 27% of adults with NAFLD during this interval [12].
While diagnosing NASH at an early stage remains a chal-
lenge, multiplexed protease-activated nanosensors have
demonstrated utility in monitoring NASH progression and
treatment response in a 3,5-diethylcarbonyl-1,4-dihydro-
collidine model of fibrosis in mice. With further develop-
ment, these noninvasive readouts can be used to diagnose
disease.
Current state of liver therapies
In order to mitigate the clinical burden of liver disease,
several therapeutic strategies have been undertaken
(Fig. 40.1).
Extracorporeal liver support devices
Liver failure is associated with abnormal accumulation of
numerous endogenous substances such as bilirubin,
ammonia, free fatty acids, and proinflammatory cytokines
[13]. Extracorporeal liver support devices have been
developed to detoxify the blood and plasma in order to
bridge patients to liver transplantation (LT) or allow the
native liver to recover from injury. Artificial liver (AL)
devices use nonliving components for detoxification, such
as membrane separation or sorbents, to selectively remove
toxins but have limited clinical use because they do not
replace the synthetic and metabolic roles of the liver [13].
Bioartificial liver (BAL) devices, on the other hand, con-
tain a cell-housing bioreactor that aims to provide the
detoxification and synthetic functions of the liver and are
an ongoing topic of clinical investigation. Current ver-
sions are either based on hollow fiber cartridges [14�16]
or on perfused three-dimensional (3D) matrices [17].
BALs, just like other hepatocyte-based therapies, face
many challenges, such as the lack of readily available
functional cell sources and the loss of cell viability and
phenotype during the treatment process.
Biopharmaceuticals
In the setting of ALF, N-acetyl cysteine (NAC) is FDA-
approved to reduce the extent of liver injury after acet-
aminophen overdose [18]. In the setting of chronic liver
disease, however, most of the FDA-approved therapies
are for hepatitis A, B, and C. A detailed listing can be
found in Table 40.1. While a few treatments have shown
moderate efficacy, there are currently no biopharmaceuti-
cals that are approved for NAFLD, NASH, or cirrhosis.
Glitazones, for example, upregulate adiponectin, an adi-
pokine with antisteatogenic and insulin-sensitizing proper-
ties [19]. Vitamin E, an antioxidant, can prevent liver
injury by blocking apoptotic pathways and protecting
against oxidative stress [19]. Despite clinical studies of a
large number of therapeutic candidates, no single agent or
combination has shown improvement to liver-related mor-
bidity and mortality in patients with NASH. Until a drug
is FDA-approved for NASH indications, lifestyle modifi-
cations and optimizing metabolic risk factors are the best
medical-treatment options for these patients.
Liver transplantation
The first attempt at human LT took place at the University
of Colorado on March 1, 1963 but turned out to be unsuc-
cessful [20]. Based on the pioneering work of Thomas
Starzl, the first extended survival of a human recipient after
LT was achieved on July 23, 1967 with a 19-month-old
female patient with hepatocellular carcinoma who survived
FIGURE 40.1 Cell-based therapies for liver disease. Extracorporeal
devices perfuse patient’s blood or plasma through bioreactors.
Hepatocytes are transplanted directly or implanted onto scaffolds.
Genetically modified large animals can be used for xenotranplantation.
738 PART | ELEVEN Gastrointestinal system
13 months before succumbing to metastatic disease [21].
After the initial success of the surgery, advancements were
made to improve donor organ quality, recipient selection,
operative and perioperative management, immunosuppres-
sion and infectious complications. These advancements have
made orthotopic LT the primary treatment for end-stage
liver disease and certain cancers. These transplants have
1-year patient survival rates over 80% [22]. However, many
challenges remain, including donor organ shortages,
recipients with more advanced disease at transplant, a grow-
ing need for retransplantation, and adverse effects associated
with long-term immunosuppression. To overcome a growing
imbalance between the supply and demand of donor livers,
transplant centers have developed strategies to expand the
donor pool. These strategies include live donor LT [23],
split-LT [24], and extended criteria for donor livers [25].
Despite all these efforts, the number of liver transplants has
not increased in the last decade.
TABLE 40.1 List of FDA-approved therapies for chronic liver diseases.
Drug
name
Year
approved
Indication(s) Mechanism
Heplisav-B 2017 Hepatitis B Combines hepatitis B surface antigen with a proprietary Toll-likereceptor 9 agonist to enhance the immune response
Mavyret 2017 HCV genotype 1�6 Fixed-dose combination of glecaprevir, an HCV NS3/4A proteaseinhibitor, and pibrentasvir, an HCV NS5A inhibitor
Vosevi 2017 Hepatitis C Fixed-dose combination of sofosbuvir, an HCV nucleotide analogNS5B polymerase inhibitor, velpatasvir, an HCV NS5A inhibitor, andvoxilaprevir, an HCV NS3/4A protease inhibitor
Ocaliva 2016 Primary biliary cholangitis FXR agonist
Zepatier 2016 HCV genotype 1 or 4 Fixed-dose combination product containing elbasvir, an HCV NS5Ainhibitor, and grazoprevir, an HCV NS3/4A protease inhibitor
Cholbam 2015 Bile acid synthesis andperoxisomal disorders
Primary bile acid synthesized from cholesterol in the liver
Daklinza 2015 HCV genotype 3 Inhibitor of NS5A, a nonstructural protein encoded by HCV
Technivie 2015 HCV genotype 4 Fixed-dose combination of ombitasvir, an HCV NS5A inhibitor,paritaprevir, an HCV NS3/4A protease inhibitor, and ritonavir, aCYP3A inhibitor
Olysio 2013 Hepatitis C Small molecule orally active inhibitor of the NS3/4A protease of HCV
Sovaldi 2013 Hepatitis C Inhibitor of the HCV NS5B RNA-dependent RNA polymerase
Incivek 2011 HCV genotype 1 Inhibitor of the HCV NS3/4A serine protease
Victrelis 2011 HCV genotype 1 Inhibitor of the HCV NS3 serine protease
Viread 2008 Hepatitis B Oral nucleotide analogue DNA polymerase inhibitor
Tyzeka 2006 Hepatitis B Inhibitor of HBV DNA polymerase
Baraclude 2005 Chronic hepatitis B withevidence of active viralreplication
Small-molecule guanosine nucleoside analog with selective activityagainst HBV polymerase
Hepsera 2002 Chronic hepatitis B withevidence of active viralreplication
Inhibitor of HBV DNA polymerase
Pegasys 2002 Chronic hepatitis C withcompensated liver disease
Binds to and activates human type 1 interferon receptors
Peg-intron 2001 Chronic hepatitis C Binds to and activates human type 1 interferon receptors
Ribavirin 2001 Chronic hepatitis C Synthetic nucleoside analog with antiviral activity
Twinrix 2001 Hepatitis A and B Recombinant vaccine
FXR, Farnesoid X receptor; HBV, hepatitis B virus; HCV, hepatitis C virus; NS3, nonstructural protein 3.
Hepatic tissue engineering Chapter | 40 739
An alternative to human LT is xenotransplantation,
though it has been clinically intractable due to concerns
about immunological rejection and zoonotic pathogen
transfer. With the advent of accessible genetic engineer-
ing technologies to circumvent the aforementioned chal-
lenges, the breeding efficiency of animals can be
leveraged to mass-produce tissue for human organ trans-
plants. Niu et al. applied CRISPR-Cas9 to inactivate all
62 copies of porcine endogenous retroviruses, thus paving
the way for pig-to-human transplants [26]. Relatedly,
Langin et al. genetically engineered porcine heart xeno-
grafts and demonstrated long-term pig-to-baboon orthoto-
pic transplantation [27].
Hepatocyte transplantation
Given the several drawbacks of LTs, alternative strategies
have been pursued. A potential alternative to LT is alloge-
neic hepatocyte transplantation (HT). Transplanted cells
can provide the missing or impaired hepatic function once
engrafted. Given their synthetic and metabolic capabili-
ties, mature hepatocytes are the primary candidates for
liver cell transplantations. HT offers several advantages
over LT. It is less invasive and can be performed repeat-
edly to meet metabolic requirements. Furthermore, multi-
ple patients can be treated with a single dissociated donor
tissue, and harvested cells can be cryopreserved for later
use on an as-needed basis.
The first experimental attempt of HT was done in
1976 to treat an animal model for Crigler�Najjar syn-
drome type I [28]. Along with other observations, it led to
the first transplant of autologous hepatocytes in 10
patients with liver cirrhosis in 1992 in Japan [29]. Since
then, reports have been published on more than 100
patients with liver disease treated by HT worldwide [30].
Human HT has resulted in partial correction of a number
of liver diseases, including urea cycle disorders [31], fac-
tor VII deficiency [32], glycogen storage disease type 1
[33], infantile Refsum disease [34], phenylketonuria [35],
severe infantile oxalosis [36], and ALF [37]. HT faces
several limitations: limited supply of high-quality mature
hepatocytes, freeze�thaw damage due to cryopreserva-
tion, poor cellular engraftment (estimated to be from
0.1% to 0.3% of host liver mass in mice after infusion of
3%2 5% of the total recipient liver cells) [30], and allo-
geneic rejection.
Clinically, the most widely used administration route
for HT is through the portal vein or one of its branches.
Hepatocytes traverse the sinusoidal vasculature and create
transient occlusions. The occlusions lead to vascular per-
meabilization which allows transplanted cells to reach the
liver parenchyma [35]. The number of cells that are
injected intraportally and subsequently engraft is a func-
tion of portal pressure and liver architecture. Thus other
administration routes have been explored for patients with
cirrhosis who have high portal pressures due to fibrosis.
In animal studies, hepatocytes transplanted into the
spleen proliferate for extended periods of time and display
normal hepatic function. The spleen has been shown to be
well-suited for hepatocyte engraftment because it func-
tions as a vascular filter and provides an immediate blood
supply [30]. The peritoneal cavity represents an attractive
administration route as it is easily accessible and can
house a large number of cells. Due to cell number
requirements associated with metabolic compensation, it
has been used in patients with ALF [38]. As an alternative
to the portal vein, spleen and peritoneum, the lymph node
(LN) has also been shown to demonstrate engraftment of
donor hepatocytes [39,40]. While this strategy has not
been utilized in the clinic yet, the preclinical data is
promising.
Current clinical trials
Several pathways have been implicated in the biology and
pathogenesis of NAFLD development: insulin resistance,
lipotoxicity, oxidative stress, altered immune/cytokine/
mitochondrial functioning, and apoptosis. New therapeu-
tic modalities are being developed to target many of these
pathways. For a detailed overview of NAFLD-targeted
drugs that are currently in the clinical trial pipeline, please
refer to Younossi et al. [41].
In vitro models
To build high-fidelity cellular models and therapies, com-
ponents of the native liver microenvironment must be
incorporated (Fig. 40.2). The liver’s highly organized
structure is key to its role as a complex tissue supporting
myriad synthetic and metabolic functions. In addition to
hepatocytes the main parenchyma of the liver, there are
several nonparenchymal cell types such as liver sinusoidal
endothelial cells, Kupffer cells, cholangiocytes, and stel-
late cells. In each lobule of the liver an array of parallel
hepatocyte cords are sandwiched between the sinusoid,
carrying circulating blood, and the bile duct, carrying
hepatocyte-secreted bile acids. Notably, this arrangement
dictates a unique set of architecturally driven cell�cell
and cell�matrix cues, which gives rise to liver-specific
phenotypes. Gradients of physicochemical stimuli along
the sinusoid drive zonal phenotypes with disparate meta-
bolic and synthetic functional profiles [42]. Interrupting
the natural order of cell arrangement in the liver is
directly connected with diseases discussed in the “Liver
disease burden” section. In this chapter, we will primarily
focus on human platforms, which are biologically distinct
from animal-derived cell models of the liver that are
reviewed more in depth elsewhere [43].
740 PART | ELEVEN Gastrointestinal system
Two-dimensional liver culture
Hepatocytes are responsible for more than 500 metabolic
and synthetic functions of the human body, often catego-
rized broadly as protein synthesis and secretion, detoxifi-
cation, bile synthesis, and nitrogen metabolism. Primary
hepatocytes quickly lose their phenotype and function
after a few days in traditional monolayer culture and
require a collagen-coated surface for adherence and sur-
vival [6]. In contrast, when primary human hepatocytes
(PHHs) are cultured between two layers of collagen gel
(i.e., sandwich culture configuration), they retain viability,
polarity and many axes of relevant metabolic and syn-
thetic function [44,45]. Guguen-Guillouzo et al. found
that a random coculture with a liver epithelial cell line
was sufficient to support hepatic albumin secretion, sug-
gesting the importance of heterotypic cell interactions for
long-term ex vivo culture [46]. Bale et al. demonstrated
that other nonparenchymal liver cells can better recapitu-
late hepatic response to inflammatory stimuli, through
higher order intercellular interactions captured in a multi-
cellular platform [47]. It was later discovered that a
micropatterned architecture consisting of hepatocyte-filled
islands surrounded by a nonbiomimetic cell type and
mouse fibroblasts can also stabilize hepatocytes, suggest-
ing the existence of conserved coculture signals across
species. These micropatterned cocultures (MPCCs) enable
the study of DILI and hepatotropic pathogen infection for
several weeks in vitro [48�50]. Furthermore, Davidson
et al. added hepatic stellate cells to the traditional MPCC
to create an in vitro model of NASH [51].
Despite the utility of two-dimensional (2D) hepatic
cultures in screening assays, a wealth of literature sug-
gests that they are dissimilar to hepatocytes in vivo.
Specifically, 2D formats, even with overlaid collagen
matrix, are more flattened than their native cuboidal
architecture. On a subcellular level, this translates to
major differences in cytoarchitecture, which is linked to
aberrant polarization and nonphysiological behavior
[52,53]. Griffith and Swartz have described the improved
presentation of relevant biochemical and mechanical cues
in 3D cultures, typically cell-laden hydrogels, compared
to traditional 2D cultures [54].
Three-dimensional liver constructs
Commonly, 3D hepatic cultures consist of primary cell or
induced pluripotent stem cell (iPSC)-derived spheroid and
organoid cultures, which are typically embedded in ECM-
based hydrogels [55]. Spheroids and organoids can be
manufactured using a variety of techniques, such as
microwell mold-based technologies, which offer a high
degree of composition and size control but are difficult to
scale. Spinning flasks and bioreactors can produce large
populations of spheroids, though they are typically non-
uniform in size and function. Bell et al. showed that pri-
mary human hepatic spheroids fabricated and cultured in
microwell plates serve as useful models of hepatotoxicity
and multiple liver pathologies [56].
Toward transplantable cell therapies, Stevens et al.
used microwell molds to create 3D hepatic spheroid
FIGURE 40.2 Advances in hepatic tissue engi-
neering. Traditional tissue culture approaches such as
the addition of extracellular matrix, soluble factors,
cocultivation with supporting cell types, hanging
drop, microwell molding, and nonadhesive surfaces
have enabled the early study of hepatocyte phenotype
in vitro in both 2D and 3D cultures. The advent of
technologies from disciplines such as chemical engi-
neering and electrical engineering has led to a new
level of control for hepatic tissue cultures, such as
micropatterning to template cell interactions, micro-
physiological systems to study the impact of bioac-
tive perfusate, polymeric biomaterials for
constructing 3D cell-laden grafts, perfusion technolo-
gies for decellularization/recellularization strategies,
and 3D printing for scalable engineering of cellular
grafts. 2D, Two-dimensional; 3D, three-dimensional.
Hepatic tissue engineering Chapter | 40 741
cocultures of PHHs and fibroblasts, which can be embed-
ded in agarose, fibrin, or polyethylene glycol hydrogel
scaffolds. The resulting tissue constructs support hepatic
function in vitro and in vivo after ectopic transplantation
into the peritoneal cavity [57]. Furthermore, Stevens et al.
demonstrated that implanting a tissue seed, consisting
of hepatic aggregates and vascular cords, into an
FRNG [fumarylacetoacetate hydrolase-deficient (Fah2/2),
recombinase activating gene-deficient (Rag12/2), nonob-
ese diabetes (NOD), and interleukin-2 receptor γchain-deficient (Il2rγ-null)] mouse model of hereditary
tyrosinemia leads to a 50-fold expansion in serum human
albumin and formation of perfusable vessels after 80 days
of cycled exposure to regenerative stimuli [58]. Relatedly,
Takebe et al. constructed liver buds consisting of iPSC-
derived hepatocyte-like cells (HLCs), mesenchymal stem
cells, and human umbilical vein endothelial cells; mesen-
teric transplantation of these human liver buds resulted in
vascularization and rescued a lethal TK-NOG mouse
model of liver injury [59].
To create clinically viable engineered liver constructs,
cell sourcing (discussed further in the “Cell sourcing” sec-
tion), and clinical-scale manufacturing of organoids and
spheroids must be addressed. Toward this end, Takebe
et al. constructed a large-scale liver bud microwell culture
platform, enabling the formation of 108 liver buds [60].
Physiological microfluidic models of liver
Despite improvements to in vitro liver platforms as model
systems and implants, static cultures lack physiologically
relevant dynamic components. The native liver’s dynamic
physiology arises from blood circulation and multiorgan
cross talk. Thus to improve biological fidelity, many have
attempted to create microfluidic models of the liver [61]. By
leveraging techniques from the semiconductor manufactur-
ing industry such as soft lithography, groups have fabricated
microphysiological systems with preformed channels to
allow for perfusion of nutrients and to aid in waste removal
[62�67]. These so-called liver-on-chip platforms allow for
the study of biochemical and mechanical cues such as
growth factor gradients and shear stress. Lee et al. demon-
strated that the presence of human stellate cells and applica-
tion of shear via flow enabled the formation of stable hepatic
spheroids on chip [68]. Furthermore, by linking microfluidic
channels between multiple tissue models, multiorgan phe-
nomena such as drug metabolite toxicity and disease pro-
gression can be captured in vitro, which is not possible in
traditional cultures [69�71].
Controlling three-dimensional architecture and
cellular organization
Another approach to improving the functionality of
tissue-engineered constructs is to more closely mimic
in vivo microarchitecture by generating scaffolds with a
highly defined material and cellular architecture, which
would provide better control over the 3D environment at
the microscale.
A range of rapid prototyping and patterning strategies
have been developed for polymers using multiple modes
of assembly, including fabrication using heat, light, adhe-
sives, or molding, and these techniques have been exten-
sively reviewed elsewhere [72]. For example, 3D printing
with adhesives combined with particulate leaching has
been utilized to generate porous poly(lactic-co-glycolic
acid (PLGA) scaffolds for hepatocyte attachment [73],
and microstructured ceramic [74] and silicon scaffolds
[75,76] have been proposed as platforms for hepatocyte
culture. Furthermore, molding and microsyringe deposi-
tion have been demonstrated to be robust methods for fab-
ricating specified 3D PLGA structures toward the
integration into implantable systems [77].
Microfabrication techniques have similarly been
employed for the generation of patterned cellular hydrogel
constructs. For instance, microfluidic molding has been
used to form biological gels containing cells into various
patterns [78]. In addition, syringe deposition in conjunc-
tion with micropositioning was recently illustrated as a
means to generate patterned gelatin hydrogels containing
hepatocytes [79]. Patterning of synthetic hydrogel systems
has also recently been explored. Specifically, the photopo-
lymerization property of poly(ethylene glycol) (PEG)
hydrogels enables the adaptation of photolithographic
techniques to generate patterned hydrogel networks. In
this process, patterned masks printed on transparencies
act to localize the ultraviolet exposure of the prepolymer
solution and, thus, dictate the structure of the resultant
hydrogel. The major advantages of photolithography-
based techniques for patterning of hydrogel structures are
its simplicity and flexibility. Photopatterning has been
employed to surface pattern biological factors [80], pro-
duce hydrogel structures with a range of sizes and shapes
[81,82], as well as build multilayer cellular networks
[83,84]. Consequently, hydrogel photopatterning technol-
ogy is ideally suited for the regulation of scaffold archi-
tecture at the multiple length scales required for
implantable hepatocellular constructs. As a demonstration
of these capabilities, photopatterning of PEG hydrogels
was utilized to generate hepatocyte and fibroblast cocul-
ture hydrogels with a defined 3D branched network,
resulting in improved hepatocyte viability and functions
under perfusion [85]. More recently, a “bottom-up”
approach for fabricating multicellular tissue constructs
utilizing DNA-templated assembly of 3D cell-laden
hydrogel microtissues demonstrates robust patterning of
cellular hydrogel constructs containing numerous cell
types [86]. Also, the additional combination of photopat-
terning with dielectrophoresis-mediated cell patterning
enabled the construction of hepatocellular hydrogel
742 PART | ELEVEN Gastrointestinal system
structures organized at the cellular scale. Overall, the abil-
ity to dictate scaffold architecture coupled with other
advances in scaffold material properties, chemistries, and
the incorporation of bioactive elements (discussed further
in the “Extracellular matrix for cell therapies” section)
will serve as the foundation for the future development of
improved tissue-engineered liver constructs that can be
customized spatially, physically, and chemically.
In vivo models
While there have been impressive advances in cell culture
models of the human liver, experimental animal models
still play an important role in the effort to engineer liver
therapies. Commonly performed surgeries such as bile
duct ligation and partial hepatectomy are experimentally
tractable models of acute liver injury, yet they are of little
clinical relevance. Drug-induced (e.g., carbon tetrachlo-
ride, acetaminophen, or thioacetamide) hepatotoxicity to
induce necrotic lesions is more recapitulative of human
pathophysiology, but the phenotype is difficult to repro-
duce [87]. In addition, modeling chronic liver injury in
animal models is problematic because they tend to rapidly
correct severe hepatic damage after a few days, which is
not representative of human disease progression and reso-
lution [88].
In order to model a human-like context for liver inju-
ries, it is necessary to develop improved, controlled mod-
els of human liver injury. Such animal models can be
useful for the evaluation of human liver biology and the
preclinical performance of candidate therapies and drugs
[89,90]. To accomplish this, human hepatocytes can be
transplanted orthotopically in immunocompromised mice
with no liver injury via the injection of a cell solution and
are useful for modeling human-specific drug metabolism,
liver injury, and hepatotropic infections. However, on
average, HT exhibits poor levels of engraftment
(B10%�30%) [91].
Transplanted hepatocytes have the ability to expand
preferentially if the host is compromised by injury or
genetic modification. The first genetically engineered
mouse model to demonstrate this was the Alb-uPA mouse
that carries a uroplasminogen activator under an albumin
promoter, causing liver injury and failure [92]. Aiming to
improve on the Alb-uPA system, a transgenic model of
hereditary tyrosinemia I was developed, in which a
genetic knockout of FAH leads to the hepatotoxic accu-
mulation of fumarylacetoacetate [93]. FAH knockout
mouse injury initiation and duration can be controlled
through the administration of a small molecule drug
[2-(2-nitro-3-trifluoro-methylbenzoyl)-1,3-cyclohexane-
dione] in the drinking water. Another inducible model
called TK-NOG, which expresses thymidine kinase
under an albumin promoter, causes hepatocyte ablation
following activation by ganciclovir treatment [94]. In
the AFC8 injury model, induction by the small mole-
cule AP20187 drives caspase-8-initiated apoptosis of
hepatocytes modified to express FK506 under an albu-
min promoter [95]. For all of the above models,
engraftment rates surpassing 70% have been observed.
A classical study in parabiotic rats in the 1960s
revealed that hepatic injury results in the expression of
systemic, soluble signals that have the potential to drive
liver regeneration [96]. Despite decades of research, the
complex signaling cascade driving liver regeneration is
still not well understood but has found utility in ectopic
humanized mouse models. Chronic liver injury often pre-
sents with high portal pressures, which can reduce
engraftment levels during an orthotopic transplantation.
Thus transplantation to ectopic sites is clinically attrac-
tive. Ectopic grafts that anastomose to the host vascula-
ture can interact with regenerative stimuli from the host
liver, causing expansion and proliferation of the trans-
planted human hepatocytes. Initially, ectopic transplanta-
tion was demonstrated in the LN [40] and spleen [97],
and later in the subcutaneous space and mesenteric fat
pad, both of which offer ease of accessibility for manipu-
lation and noninvasive imaging [58,59,98].
Taken together, the range of liver injury animal mod-
els are an essential tool for studying various perturbations
to normal liver biology and building implantable tissue
constructs to address acute and chronic liver failure. The
field is just beginning to uncover mechanisms that control
liver regeneration in various disease and injury contexts.
The discovery of new soluble regenerative signals will be
central to advancing therapies that have the potential to
improve the supply of donor tissue.
Cell sourcing
Cell number requirements
The development of cell-based therapies poses myriad
challenges, partially stemming from the scale of the liver.
An adult human liver is estimated to possess 241 billion
hepatocytes, 24 billion stellate cells, and 96 billion
Kupffer cells [99]. Sourcing such enormous cell numbers
using current technologies is not feasible.
However, many human HT studies suggest that clini-
cal intervention is possible with a fewer number of cells
and offer critical insights to help us determine minimum
cell numbers. In a review, Fisher and Strom cataloged 78
different human HT studies, detailing both the input cell
number and a qualitative description of functionality [30].
Correlating these, we can broadly surmise that to correct
inborn errors of metabolism, at least 1�10 billion hepato-
cytes are needed. However, for ALF, that number grows
to 5�20 billion cells. For liver cirrhosis, HTs have largely
been unsuccessful (discussed further in the “Hepatocyte
transplantation” section); therefore it is unclear what the
Hepatic tissue engineering Chapter | 40 743
cellular requirements for cirrhosis are. While injection of
hepatocytes is not the same as implantation within a scaf-
fold, these studies serve as useful inputs into more com-
plex physiological models.
In order to get us closer to these numbers, many dif-
ferent cell sources have been explored.
Immortalized cell lines
Immortalized hepatocyte cell lines can be derived from
liver tumor tissue or directly from primary hepatocytes
in vitro. The prominent lines utilized today are HepG2,
derived from hepatocellular carcinoma; HepaRG [100], a
human bipotential progenitor cell line; C3A, derived from
HepG2s; and Huh7, derived from liver tumor [101].
Several other fetal and adult hepatic cell lines have also
been established, typically using a combination of viral
oncogenes and the human telomerase reverse transcriptase
protein [102]. However, these cell lines lack the full func-
tional capacity of primary adult hepatocytes and there is a
risk that oncogenic factors could be transmitted to the
patient, limiting their use as a cell source for transplanta-
tion therapies.
Primary cells
Unlike immortalized lines, PHHs can provide a whole
host of human liver-specific function. PHHs, however, are
limited in supply, and their phenotype is difficult to main-
tain in vitro. Many methods have been developed for
maintaining long-term functionality of hepatocytes
through the use of a variety of configurations and bioma-
terial constructs, which are further discussed in the “In
vitro models” section. Due to limitations in the supply of
mature human hepatocytes, many groups have attempted
to promote the expansion and proliferation of PHHs
in vitro. Peng et al. have shown that TNFα promotes the
expansion of hepatocytes in 3D cultures and enables serial
passaging and long-term culture for more than 6 months
[103]. In a similarly notable study, Hu et al. identified an
optimal cell culture cocktail consisting of B27 supplement
(without vitamin A), R-spondin, CHIR99021 (a Wnt ago-
nist), NAC, nicotinamide, gastrin, epidermal growth fac-
tor (EGF), TGFα, fibroblast growth factor (FGF)7,
FGF10, HGF, a TGFβ inhibitor (A83-01), and ROCK
inhibitor that led to long-term 3D organoid culture of
PHHs [104].
Fetal and adult progenitors
Given their ability to differentiate into diverse lineages
both in vitro and in vivo, iPSC and human embryonic
stem cell cultures can also be utilized to generate HLCs.
Various differentiation protocols have been applied to
these cultures to yield cell populations that exhibit some
phenotypic and functional characteristics of hepatocytes
[62,105�108]. These populations are termed HLCs
because of their expression of fetal proteins and fetal-like
cytochrome P450 profiles [109]. While they are distinct
from mature adult hepatocytes, HLCs can still serve as a
potential cell source in very specific contexts.
In addition to pluripotent cells, bipotential progenitor
cells can also serve as a source for hepatocytes. Huch
et al. delineated conditions that allow for long-term
expansion of adult bile duct-derived EpCAM1 bipoten-
tial progenitor cells from the human liver [110]. The
expanded cell population attained using their protocol is
stable at the chromosomal level and can be converted into
functional HLCs in vitro and in vivo [110].
Reprogrammed hepatocytes
HLCs can also be generated using direct reprogramming
of mature cell types. For example, several groups have
demonstrated the feasibility of reprogramming fibroblasts
into HLCs without a pluripotent intermediate [111�113].
Cheng et al. demonstrated that a combination of nuclear
factors can stimulate the conversion of hepatoma cells to
HLCs [101]. These findings raise the future possibility of
deriving human HLCs directly from another adult cell
type.
Extracellular matrix for cell therapies
The ECM of the liver provides a structural scaffold with
bioactive cues that modulate hepatic function and pro-
mote vascularization. Collagen and fibronectin are the
major structural components of the liver ECM. Along
with other nonstructural proteins, these components par-
ticipate in integrin-mediated signaling between cells and
their surrounding matrix. Hepatocytes are sensitive to
their ECM, and it has been demonstrated that the presence
of abnormal amounts and/or types of ECM components
correlates with the onset and progression of liver fibrosis
[114].
ECM scaffolds for hepatic tissue engineering are use-
ful for constructing 3D tissue models and as a delivery
vehicle for implants. Polymeric biomaterial hydrogels
gained popularity as an engineering tool for recapitulating
a physiologically relevant 3D tissue niche. Aside from
creating a permissive environment for hepatocyte survival
and growth, ECM scaffolds for hepatic tissue engineering
also enable the formation of biliary and vascular networks
that will be further discussed in the “Vascular and biliary
tissue engineering” section. Broadly speaking, ECM scaf-
folds can be constructed using synthetic and/or naturally
derived polymers.
744 PART | ELEVEN Gastrointestinal system
Natural scaffold chemistry and modifications
A wide range of natural biomaterial polymers spanning
polysaccharides (e.g., dextran and chitosan), peptides
(e.g., collagen and fibrin), decellularized ECM (dECM),
and composites of these have been employed as hepatic
tissue�engineering scaffolds [55,57,58,115�121]. The
advantages of biologically derived materials include their
biocompatibility; naturally occurring cell adhesive moie-
ties; and, in the case of decellularization, native architec-
tural presentation of ECM molecules. However, naturally
derived biomaterials have several barriers to use in the
clinic, primarily due to lot-to-lot variability and xenoge-
neic origin.
The choice of material determines the physicochemi-
cal and biological properties of the scaffold. For
example, early efforts in developing implantable hepatic
constructs utilized collagen-coated dextran microcarriers
that enabled hepatocyte attachment since hepatocytes
are known to be anchorage-dependent cells. The intra-
peritoneal transplantation of these hepatocyte-attached
microcarriers resulted in successful replacement of
liver functions in two different rodent models of genetic
liver disorders [122]. Subsequently, collagen-coated or
peptide-modified cellulose [120,123], gelatin [124], and
gelatin�chitosan composite [125] microcarrier chemis-
tries have also been explored for their capacity to promote
hepatocyte attachment. On the other hand, materials that
are poorly cell adhesive such as alginate [115] have been
exploited for their utility in promoting hepatocy-
te�hepatocyte aggregation (i.e., spheroid formation) and
phenotypic stabilization within these scaffolds.
Collectively, the size of engineered tissues created by
these approaches is limited by oxygen and nutrient diffu-
sion to only a few hundred microns in thickness.
To address this constraint, recent work has sought to
use decellularized whole organ tissue as a matrix for liver
tissue engineering. The decellularization process utilizes
perfusion-based technologies to remove cells from donor
tissues but preserve the structural and functional charac-
teristics of the native underlying tissue. Recent advances
in decellularization protocols have yielded scaffolds with
native liver ECM composition, growth factor presentation,
vascular structure, and biliary network architecture
[126�128]. To date, seeding protocols have achieved up
to 95% efficiency of recellularization with relevant cell
populations (e.g., hepatocytes, vascular cells and bipotent
hepatic progenitors); resulting recellularized grafts exhib-
ited liver-specific function, and survival after transplanta-
tion in rodents [121,126,129]. Furthermore, cell-laden,
xeno-derived dECM scaffolds are compatible with immu-
nocompetent animal models [128]. However, given the
shortage of donor tissue, the wide use of dECM scaffolds
is unlikely.
Synthetic scaffold chemistry
In contrast to biologically derived material systems, syn-
thetic materials enable precisely customized architecture
(porosity and topography), mechanical and chemical
properties, and degradation modality and kinetics, which
are known to drive cell behavior. Synthetic materials that
have been explored for liver tissue engineering include
poly(L-lactic acid) (PLLA), PLGA, poly(ε-caprolactone),and PEG [85,98,130�135]. Polyesters such as PLLA and
PLGA are the most common synthetic polymers utilized
in the generation of porous tissue-engineering constructs.
These materials are biocompatible, biodegradable, and
have been used as scaffolds for HT [132,136]. A key
advantage of PLGA is the potential to finely tune its deg-
radation time due to differences in susceptibility to hydro-
lysis of the ester groups of its monomeric components
(lactic acid and glycolic acid). However, the accumulation
of hydrolytic degradation products has been shown to pro-
duce an acidic environment within the scaffold which
initiates peptide degradation and stimulates inflammation,
which may affect hepatocyte function [137].
Consequently, as alternatives to macroporous scaffold
systems, approaches aimed at the efficient and homoge-
neous encapsulation of hepatocytes within a fully 3D
structure have been explored. In particular, hydrogels that
exhibit high water content and thus similar mechanical
properties to tissues are widely utilized for various tissue-
engineering applications, including hepatocellular plat-
forms. Synthetic, PEG-based hydrogels have been
increasingly utilized in liver tissue-engineering applica-
tions due to their high water content, hydrophilicity, resis-
tance to protein adsorption, biocompatibility, ease of
chemical modification, and the ability to be polymerized
in the presence of cells, thereby enabling the fabrication
of 3D networks with uniform cellular distribution [138].
PEG-based hydrogels have been used for the encapsula-
tion of diverse cell types, including immortalized and pri-
mary hepatocytes and hepatoblastoma cell lines
[85,98,135]. The encapsulation of primary hepatocytes
requires distinct material modifications [e.g., 10% w/v
PEG hydrogel, inclusion of RGD adhesive motifs] as
detailed below, as well as, analogous to 2D coculture sys-
tems, the inclusion of nonparenchymal supporting cell
types such as fibroblasts and endothelial cells [135].
Modifications in scaffold chemistry
The relatively inert nature of synthetic scaffolds allows
for the controlled incorporation of chemical/polymer moi-
eties or biologically active factors to regulate different
aspects of cellular function. Chemical modifications such
as oxygen plasma treatment or alkali hydrolysis of PLGA
[139,140], or the incorporation of polymers such as
Hepatic tissue engineering Chapter | 40 745
poly(vinyl alcohol) or poly(N-p-vinylbenzyl-4-O-β-D-galactopyranosyl-D-glucoamide) (PVLA) into PLGA or
PLLA scaffolds [132,141,142] have improved hepatocyte
adhesion by modulating the hydrophilicity of the scaffold
surface [143]. Biological factors may include whole bio-
molecules or short bioactive peptides. Whole biomole-
cules are typically incorporated by nonspecific adsorption
of ECM molecules such as collagen, laminin, or fibronec-
tin [139,144]; covalent conjugation of sugar molecules
such as heparin [145,146], galactose [131,147], lactose
[145] or fructose [148]; or growth factors such as EGF
[149]. Alternatively, short bioadhesive peptides that inter-
act with cell surface integrin receptors have been exten-
sively utilized to promote hepatocyte attachment in
synthetic scaffolds. For example, conjugation of the RGD
peptide to PLLA has been shown to enhance hepatocyte
attachment [150], whereas RGD modification signifi-
cantly improved the stability of long-term hepatocyte
function in PEG hydrogels [98,135]. The additional incor-
poration of adhesive peptides that bind other integrins
may serve as a way to further modulate and enhance
hepatocyte function within synthetic polymer substrates.
Moreover, Stevens et al. demonstrated that integration of
matrix metalloproteinase-sensitive peptide sequences into
hydrogel networks as degradable linkages has been shown
to enable cell-mediated remodeling of the hydrogel [151].
The capacity to modify biomaterial scaffold chemistry
through the introduction of biologically active factors will
likely enable the finely tuned regulation of cell function
and interactions with host tissues important for
implantable systems.
Porosity
A common feature of many implantable tissue-
engineering approaches is the use of porous scaffolds that
provide mechanical support, often in conjunction with
cues for growth and morphogenesis. Collagen sponges,
various alginate and chitosan composites, and PLGA are
the most commonly used porous scaffolds for hepatocyte
culture and are generally synthesized using freeze-dry or
gas-foaming techniques. Pore size has been found to regu-
late cell spreading and cell�cell interactions, both of
which can influence hepatocyte functions [116], and may
also influence angiogenesis and tissue ingrowth [152].
Porous, acellular scaffolds are normally seeded using
gravity or centrifugal forces, capillary action, convective
flow, or through cellular recruitment with chemokines,
but hepatocyte seeding is generally heterogeneous in these
scaffolds [153,154].
Vascular and biliary tissue engineering
Beyond compatibility with hepatic cell types, scaffolds
should also be conducive to vasculature formation.
Relying on vascularization by the host is not sufficient for
large tissue constructs required for the clinic, because
cells that are not near capillary structures
(. 150�200 μm) are at a risk for necrosis after a matter
of hours due to a lack of oxygen, nutrient availability, and
waste transport. In this section, we discuss composite
approaches toward building scalable, vascularized
constructs.
Vascular engineering
Approaches to engineering vessels can generally be cate-
gorized as bottom-up induction of vascular assembly and
top-down fabrication of vascular conduits [155]. Bottom-
up vascular engineering approaches are built upon the
idea of neovascularization, or new vessel formation.
Vessel formation can occur by angiogenic sprouting, the
formation of vessels branching off of an existing blood
vessel, or vasculogenesis, the self-assembly of single
endothelial cells or progenitor cells into lumenized ves-
sels. Despite the ability of single vascular cells to coa-
lesce to enable self-assembly, vasculogenesis is
accelerated by coculture with supporting stromal cells,
such as fibroblasts, mesenchymal stem cells, and pericytes
[155�157]. Studies exploring angiogenesis and sprouting
events suggest that chemical gradients, fluid-driven shear
stress, and a hypoxia are key players in vessel formation
[158�161].
Top-down fabrication approaches dictate geometry
and architecture, rather than driving self-assembly.
Polymer molding using microetched silicon has been
shown to generate extensive channel networks with capil-
lary dimensions, though it is not amenable to high-
throughput manufacturing [162]. While direct printing of
cells can be cytotoxic, 3D printing has become a popular
approach for fabricating hollow channels that enable vas-
cular cell seeding. The challenge lies in using this
approach to build patent capillary beds, which are
5�10 μm in diameter. While two-photon polymerization
has an impressively high feature size resolution at
100 nm, the trade-off between build volume (i.e., building
constructs large enough for clinical impact), build speed
(i.e., impacting manufacturability), and printer resolution
renders it inappropriate for most applications in tissue
engineering. Alternate printers using direct-ink writing
(DIW) of viscoelastic materials have emerged as a power-
ful tool for the fabrication of patterned hydrogel con-
structs. DIW can achieve minimum feature size
resolutions from 1 to 250 μm, depending on the ink
“building block” size [163]. However, DIW printing
requires yield-stress fluid inks with restrictive viscosities
(102�106 mPa s), such that they fluidize under stress but
regain the original shear elastic modulus after printing
[163]. Kolesky et al. demonstrated that DIW and fugitive
inks were useful for multimaterial printing as well as
746 PART | ELEVEN Gastrointestinal system
construction of thick (. 1 cm), vascularized tissues
[164,165]. Miller et al. demonstrate the use of thermal
microextrusion approach to create sacrificial sugar glass
lattices that can be embedded in various cell-laden bioma-
terials, evacuated and subsequently lined with vascular
cells [166].
Taken together, self-assembly driven formation of a
capillary bed can be combined with printing of larger ves-
sels to enable the fabrication of a fully vascularized tissue
construct. Song et al. used 3D printing to create curved
vascular channels using a fugitive ink [159]. The vascular
cells lining the channel were able to degrade the sur-
rounding hydrogel matrix and undergo sprouting when
exposed to angiogenic factors. They further demonstrated
that the curvature and complexity of printed vasculature
impacted the extent of sprouting, which will be an impor-
tant consideration for vessel patterning in future clinical
applications [159].
Host integration
For vascularized constructs to survive after implantation
the engineered vessels must anastomose with the host vas-
culature. To date, the exact mechanism that drives inte-
gration with the host is not well understood [155].
A number of studies have elucidated the contribution of
cytokines important in angiogenesis and recruitment of
host vasculature in implant constructs, such as vascular
EGF (VEGF) [136], basic FGF [167], and VEGF in com-
bination with platelet-derived growth factor [168].
Furthermore, preimplantation of VEGF releasing alginate
scaffolds prior to hepatocyte seeding was demonstrated to
enhance capillary density and improve engraftment [169].
In addition, while building an AL tissue construct,
Stevens et al. demonstrated that the coimplantation of par-
allel endothelial cell cords with PHH and fibroblast spher-
oids led to better hepatic performance and survival when
compared to an implant with nonpatterned endothelial
cells, suggesting that there may be an optimal templated
endothelial geometry that enables vascularization and
host integration in vivo [58]. Surgical anastomosis poses
an alternative to biologically driven anastomosis, though
this approach requires invasive surgery and access to
suture-able vessels both in the graft and in the host.
Strategies to incorporate vasculature into engineered con-
structs include the microfabrication of vascular units with
accompanying surgical anastomosis during implantation
[75,170].
In addition to interactions with the vasculature, inte-
gration with other aspects of host tissue will constitute
important future design parameters. For instance, incorpo-
ration of hydrolytic or protease-sensitive domains into
hepatocellular hydrogel constructs could enable the degra-
dation of these systems following implantation [151]. Of
note, liver regeneration proceeds in conjunction with a
distinctive array of remodeling processes such as protease
expression and ECM deposition. Interfacing with these
features could provide a mechanism for the efficient inte-
gration of implantable constructs. Similarly to whole liver
or cell transplantation, the host immune response follow-
ing the transplant of tissue-engineered constructs is also a
major consideration. Immunosuppressive treatments will
likely play an important role in initial therapies, although
stem cell�based approaches hold the promise of
implantable systems with autologous cells. Furthermore,
harnessing the liver’s unique ability to induce antigen-
specific tolerance [171] could potentially represent
another means for improving the acceptance of engi-
neered grafts.
Biliary network engineering
Importantly, future iterations of hepatic grafts should
include a biliary system, which is responsible for excre-
tory functions. In a similar vein, we envision that a com-
bination of “top-down” manufacturing, such as the
aforementioned technologies for generating patent vascu-
lar conduits, and “bottom-up” approaches, which could
involve leveraging biological phenomena that drive bili-
ary morphogenesis, will be useful for building a biliary
network.
The biliary tree is a complex, 3D network of tubular
conduits of various sizes and properties. The liver con-
tains an intrahepatic compartment that is lined by biliary
epithelial cells, termed cholangiocytes, that aid in the
modification and removal of hepatocyte-secreted bile.
Even though the blood vessel�fabrication approaches
delineated above have not yet been applied to engineering
bile networks for implantable liver constructs, advance-
ments in cholangiocyte sourcing methods have made it
feasible. Sampaziotis et al. identified a protocol for
directed differentiation of human iPSCs into
cholangiocyte-like cells [172]. In 2017 the same group
provided the first proof-of-concept study to reconstruct
the gallbladder wall and repair the biliary epithelium
using human primary cholangiocytes expanded in vitro
[173]. Furthermore, current studies are focused on the
development of in vitro models that exhibits biliary mor-
phogenesis and recapitulates appropriate polarization and
bile canaliculi organization [174�176], as well as plat-
forms for the engineering of artificial bile duct structures
[177,178].
Conclusion and outlook
Traditionally, hepatic tissue�engineering research has
focused on designing the microenvironment to support a
stable hepatic phenotype. As concomitant advances in
pluripotent cell research and polymer chemistry have
been actualized, new cell sources and extracellular
Hepatic tissue engineering Chapter | 40 747
matrices have been added to the pipeline. This interdisci-
plinary synergy has been the driving force behind the
development of tissue-engineered grafts with long-term
survival and growth. In order to inch closer to the regen-
erative medicine north star of an ex vivo engineered graft
that can serve as a replacement for the native organ, there
are several areas to consider for improvement: (1) vascu-
larization of thick, dense grafts through a combination of
self-assembly and bioprinting; (2) engineering of the
hepatic graft to prevent immune rejection in allogeneic
and xenogeneic settings; (3) improved understanding of
metabolic and cellular requirements of various liver dis-
eases; (4) development of scalable, renewable cell sources
that do not compromise the functional capabilities of
cells; (5) leveraging of animal injury models as bioreac-
tors for cell sourcing; and (6) upscaling of grafts in a
manner that is compatible with FDA’s Good
Manufacturing Practice standards.
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