Towards a Synthetic Tissue Model of the Lower Urinary Tract

Alexander Preis

, Christina Merkl

, Paula Miralles

, Svenja Heer

, Elisabeth Benke

Sebastian Reitelshöfer

, Sina Martin

, Ralf Rieker

and Jörg Franke

Institute for Factory Automation and Production Systems, Friedrich-Alexander-Universität Erlangen-Nürnberg, Germany

Institute of Pathology, Universitätsklinikum Erlangen, Friedrich-Alexander-Universität Erlangen-Nürnberg, Germany

ralf.rieker@uk-erlangen.de

Keywords: Anatomical Model, Urology, Lower Urinary Tract, Urinary Bladder, Urethra, Artificial Urine.

Abstract: The development of medical devices often depends on in vivo studies to validate the proper functioning of

the products. These trials provide ethical as well as economic challenges, which can be partially addressed by

the usage of realistic synthetic tissue models that replicate human anatomy and the corresponding properties

of the biological tissue. In this work, a silicone-based model with a fiber structure and a PVA-based model

that exhibits fabrication-induced anisotropy are presented in the special context of the lower urinary tract. The

analysis of the materials in the uniaxial tensile test shows the anisotropic and viscoelastic properties of the

materials. Furthermore, the anatomical model of the lower urinary tract shows expected deformation in

simulation as well as in the real silicone model. Additionally, a suitable artificial urine according to ISO 20696

is shown for use with the model. First experiments to change the pH of the artificial urine are successfully

conducted.

1 INTRODUCTION

In the development of medical devices, a large

proportion of the costs occur due to elaborate in vivo

studies on animals and humans (J. A. DiMasi et al.,

2016). As shown in Figure 1, approximately 2.8

million animals were used for research purposes in

Germany alone in 2014 (DFG, 2021).

Often previously unknown weaknesses of the

tested medical device become apparent in the course

of the in vivo test phases, which leads to further

iterations, further animal testing and means an

economic loss for the companies (I. S. Yoo et al.,

2020). In addition, a successful animal test study does

not necessarily indicate the suitability of the product

for use in humans, as there are sometimes significant

differences both anatomically and physiologically

(M. Viceconti et al., 2016). To overcome the ethical

as well as economic challenges of animal testing,

realistic synthetic tissue models that replicate human

https://orcid.org/0000-0003-3469-5982

https://orcid.org/0000-0002-6610-4430

https://orcid.org/0000-0002-4472-0208

https://orcid.org/0000-0002-2146-8265

https://orcid.org/0000-0003-0700-2028

anatomy and the properties of the biological tissues

are an obvious choice.

Figure 1: In Germany, approximately 2.8 million animals

were used for research purposes in 2014 (DFG, 2021).

In this work, first steps towards a synthetic in-

vitro and in-silico tissue model of the lower urinary

tract are presented that will be used for the testing of

intraurethral artificial urinary spincters like the one

presented in (A. Preis et al., 2022).

190

Preis, A., Merkl, C., Miralles, P., Heer, S., Benke, E., Reitelshöfer, S., Martin, S., Rieker, R. and Franke, J.

Towards a Synthetic Tissue Model of the Lower Urinary Tract.

DOI: 10.5220/0011780500003414

In Proceedings of the 16th Inter national Joint Conference on Biomedical Engineering Systems and Technologies (BIOSTEC 2023) - Volume 1: BIODEVICES, pages 190-197

ISBN: 978-989-758-631-6; ISSN: 2184-4305

 2023 by SCITEPRESS – Science and Technology Publications, Lda. Under CC license (CC BY-NC-ND 4.0)

1.1 Structure of Biological Tissues

Biological tissues often have unique properties

compared to many engineering materials. They are

capable of load-dependent adaptation and repair

when damaged, which means, that they can adapt to

changing mechanical requirements (remodeling).

Most biological tissues are complex composite

materials with inhomogeneous and anisotropic

properties. As a result, the mechanical properties vary

from point to point within a tissue and the response to

forces acting in different loading directions can be

distinct. As an example, the values for the strength

and stiffness of bone differ both between various

bones and at individual points within the same bone.

Furthermore, biological tissues are viscoelastic to a

large extent, so the rate of load application and the

creep and relaxation processes that occur are also

relevant to the analysis of mechanical properties. (Y.

C. Fung, 1993)

The mechanical properties of biological tissues,

which are mainly composed of cells and an

extracellular matrix consisting of fibers and ground

substance, are largely determined by the collagen and

elastin fiber content. The properties of the tissues are

optimized by different compositions and orientations

according to the specific application. (Y. C. Fung,

1993)

The primary mechanical function of collagen

fibers is to resist axial tension. Because of their large

length-to-diameter ratio, they are susceptible to

buckling and not suited to withstand compressive

loads. When a fiber is pulled, its length increases, as

with a mechanical spring, and the energy applied to

stretch the fiber is stored. The release of this energy

subsequently returns the fiber to its unstretched

configuration. The individual fibrils of collagen are

surrounded by a gel-like ground substance, which is

largely composed of water and contributes to the

viscoelastic material behavior with relatively high

tensile and low compressive strength of the collagen

fibers. Among the non-collagenous tissue

components, elastin is another important fibrous

protein with material properties comparable to those

of rubber. The elastic elastin fibers are highly

extensible, and their elongation is reversible even

under high stress. They behave elastically with low

stiffness up to an elongation of about 200 %, followed

by a short range where the stiffness increases sharply

until failure. In summary, elastin fibers exhibit elastic

material properties with low Young's modulus, while

collagen fibers exhibit viscoelastic material behavior

with higher Young's modulus. (N. Özkaya et al.,

2012)

1.2 Mechanics of Biological Tissues

Viscoelastic materials like biological tissues exhibit

both elastic and viscous behavior. In the following,

elasticity, viscosity, and viscoelasticity are briefly

explained for better understanding.

Elasticity

Elasticity describes the ability of a material or body

to reverse a change in shape caused by an external

force through its own internal force, which is also

referred to as the restoring force. If a specimen with

an original length L

is loaded longitudinally with a

force that deforms the specimen to length L, the strain

ε is defined according to (1):

𝜀=

𝐿−𝐿



𝐿



∆𝐿

𝐿



(1)

with:

𝜀 strain

𝐿 current length

𝐿

initial length

∆𝐿 change in length

Stress σ and strain are related by Hooke's law (2)

and are proportional to each other. The

proportionality factor is a material constant which is

called Young's modulus 𝐸.

𝜎=𝐸∙𝜀

(2)

with:

𝜀 strain

𝐸 Young’s modulus

𝜎 stress

In models, elastic behavior is usually represented

by linear elastic springs. They can deform reversibly

when a load is applied and return to their original

shape when the load is removed. If a material behaves

linearly elastic, the stress is linearly proportional to

the strain. In elastic materials, the mechanical

properties are independent of time. When externally

loaded, they deform instantaneously and return to

their original shape almost immediately when

unloaded. The fibrous portion of biological tissues

can be simplified in such a way. (N. Özkaya et al.,

2012)

Viscosity

Viscosity provides information about the flow

behavior of a material. The dynamic viscosity η

describes the internal friction of fluids and represents

the resistance to a forced, irreversible change of

location of its volume elements. For a better

illustration, one can imagine that the fluid is located

Towards a Synthetic Tissue Model of the Lower Urinary Tract

191

between two plates oriented parallel to each other and

adheres to both plates. If the upper plate is moved

with a velocity 𝑣, the fluid layer in the immediate

vicinity also moves with the velocity 𝑣 due to

adhesion. Since the lower plate has not moved, the

adjacent fluid layer is at rest. The velocity within the

fluid increases from the lower plate at rest to the

moving upper plate. In a laminar flow, the velocity

gradient 𝑑𝑣/𝑑𝑦 arises, which is also often

abbreviated as 𝛾

 . For ideal fluids, Newton's law

applies to calculate the resulting shear stress τ

according to (3):

𝜏= 𝜂

𝑑𝑣

𝑑

𝑦

= 𝜂 𝛾

(3)

with:

𝜏 shear stress

𝜂 dynamic viscosity

𝛾



velocity gradient

The dynamic viscosity 𝜂 acts as a constant of

proportionality. Thus, in contrast to elastic solids, the

stresses do not depend on the deformation per se, but

on the rate of deformation. In models, purely viscous

behavior is therefore usually represented by a

Newtonian damper. In this model, the resulting stress

at a constant viscosity depends only on the

deformation rate and the strain rate. The ground

substances portion of biological tissues can be

simplified as purely viscous. (J. de Vicente, 2012)

Viscoelasticity

Because of their composite structure, biological

tissues exhibit both elastic and viscous properties.

They therefore exhibit creep and relaxation processes

when subjected to loading and unloading. The

response of viscoelastic materials depends on the rate

at which the load is changed. As a result, the stress-

strain diagram of a viscoelastic material also depends

on the rate at which the load is applied to the material.

Its behavior can be described by characteristic

material functions, which are determined in special

experiments, first and foremost the creep and

relaxation test. Here, the response of a material to a

constant stress applied at time 𝑡

and removed later at

time 𝑡

is observed. Such a stress immediately causes

a strain 𝜀 in a linearly elastic material at time 𝑡

. This

constant strain remains in the material until time 𝑡

If the applied stress is removed at time 𝑡

, the linear

elastic material immediately and completely recovers

from the deformation. To the same constant loading

condition, a viscoelastic material responds with strain

that gradually increases between times 𝑡

and 𝑡

. At

time 𝑡

, when the load is removed, a gradual recovery

begins.

To represent viscoelastic behavior, models that

have springs as well as dampers are used as basic

elements. One of the best-known models is the

Kelvin-Voigt model, which is based on a parallel

connection of the two basic elements. Due to the

parallel connection, the strain of the damper equals

that of the spring. Using Hooke's law for the elastic

spring and Newton's law for the damper, the first

order differential equation shown in (4) is obtained

(Y. C. Fung, 1993):

𝜎= 𝜎



+𝜎



=𝐸 𝜀+𝜂𝜀 (4)

with:

𝜎 total stress

𝜎

𝐻

stress of the spring

𝜎

𝑁

stress of the damper

𝐸 Young's modulus

𝜀 strain of the spring

𝜂 dynamic viscosity

𝜀



deformation rate

The deformation of a damper placed parallel to a

spring, as in the Kelvin-Voigt model, is limited by the

response of the spring to the applied forces. The

damper cannot deform continuously in this

arrangement. Therefore, the Kelvin-Voigt model

represents a viscoelastic solid behavior. (N. Özkaya

et al., 2012)

Another well-known model for simulating

viscoelastic behavior is the Maxwell model. Like the

Kelvin-Voigt model, it consists of a spring and a

damper. However, these are not connected in parallel,

but in series. From the arrangement it follows that the

total stress of the system must be equal to the stress

in the spring and to the stress in the damper. The total

strain of the system results from the individual strains

of the spring and the damper, as described in (5):

𝜀= 𝜀



+𝜀



=𝐸 𝜎+𝜂𝜎 =𝐸𝜂𝜀

(5)

with:

𝜀 total strain

𝜀

𝐻

strain of the spring

𝜀

𝑁

strain of the damper

𝐸 Young's modulus

𝜎 total stress

𝜂 dynamic viscosity

𝜀



deformation rate

In the case of the Maxwell model, a force

application leads to deformation of both the spring

and the damper. The deformation of the spring is

finite, whereas the damper deforms as long as the

force is applied to the system. Therefore, the overall

behavior of the Maxwell model resembles a fluid

BIODEVICES 2023 - 16th International Conference on Biomedical Electronics and Devices

192

rather than a solid and is referred to as a viscoelastic

fluid model. (N. Özkaya et al., 2012)

However, the Kelvin-Voigt and Maxwell models

alone are not capable of representing the real behavior

of many viscoelastic materials but can be used to

create more complex viscoelasticity models. A well-

known one is the Zener model, which is also referred

to as the standard linear solid model and can be

described by two equivalent representations: the first

consists of a series connection of a Kelvin model with

a spring, and the second one consists of a parallel

connection of a Maxwell model with a spring. It is

used to represent the viscoelastic behavior of some

biological materials, such as cartilage. Similarly,

there is also the three-element fluid model used in the

study of blood. In this one, the additional spring is

replaced by a damper. Also, other models can be

created by combining any number of Maxwell and/or

Kelvin-Voigt bodies to represent the behavior of

other materials. (N. Özkaya et al., 2012)

2 MATERIALS AND METHODS

In the present work, two possible materials for a

synthetic in-vitro tissue model of the lower urinary

tract are presented. Additionally, a suited anatomical

geometry and a recipe for artificial urine, which can

be used to test urinary stone formation, are shown.

2.1 Silicone-Based Model

For the silicone-based model, the RTV-2 silicone

"Elastosil P7670 A/B" with a 10 wt% silicone oil

(AK 100) content is used (Wacker Chemie AG). To

create the anisotropy, additional ‘fibers’ without

silicone oil are embedded in a matrix of silicone

according to the previous formula. The silicone is

mixed in a vacuum stirrer to avoid air inclusion and

subsequently cast into sheets with a thickness of

2 mm and fully cured. The resulting tissue model has

a fiber volume content of around 24 %.

2.2 PVA-Based Model

For the PVA-based model, a solution consisting of 10

wt% PVA powder (MW: 133000 g/mol, degree of

hydrolysis: 99 %, Polysiences Inc.) in distilled water

is prepared. Both components are mixed in a sealed

vessel at a constant temperature of 100 °C for a

duration of 6 h. The finished, transparent solution is

poured into sheets of 2 mm thickness, analogously to

the silicone. Any air pockets are removed by vacuum

and the plate is cooled to room temperature.

Afterwards it is frozen for at least 16 hours. After the

first freezing cycle (consisting of freezing and

complete thawing to room temperature), the

previously viscous mass has a gelatinous consistency

and a whitish color. After two freezing cycles, two

opposite ends of the sheet are fixed in a jig, the sheet

is stretched by 80 % of its original length and eight

more freezing cycles are performed to create the

anisotropy. The material is packed airtight to keep the

samples from drying out.

2.3 Material Characterization

To determine the mechanical properties of the tissue

models, uniaxial tensile tests (Z 2.5/TN1S,

ZwickRoell GmbH & Co. KG) are performed with

the uniform base materials as well as along and

perpendicular to the fiber orientation and strain

direction of the anisotropic tissue models. For this

purpose, specimen geometry S3A of the DIN 53504

standard is used at strain rates of 200 and

800 mm/min. The significance of the results is then

statistically analyzed.

2.4 Lower Urinary Tract

The lower urinary tract consists of the urinary bladder

(vesica urinaria) and urethra. They work together as

a functional unit and perform the tasks of storing and

emptying urine. In both sexes, the vesica urinaria is

located in the lesser pelvis just behind the symphysis.

In women, it lies in front of the vagina and in front of

and below the uterus; in men, it lies in front of the

rectum. The bladder of an adult has a capacity of

about 400 to 500 ml and is emptied to less than 50 ml

during micturition. Depending on the state of filling,

it is bowl-shaped flattened or spherical. The thickness

of the bladder wall varies according to the volume of

urine, decreasing accordingly as it expands. The wall

thickness ranges from 1 to 5 mm, and can also reach

up to 10 mm. The transition from the vesica urinaria

to the urethra is called the bladder neck. The female

urethra is straight and short, the male urethra passes

through the penis and is longer and has several

curves. The female urethra considered for this work

is about 40 mm long and has a diameter of about 8

mm. (D. Schultz-Lampel et al., 2012; M. Schünke et

al., 2022)

The anatomical model of the urinary bladder was

created using Autodesk Inventor (Autodesk Inc.).

Afterwards the deformation during normal

micturition with an abdominal pressure of 20 cmH

and a detrusor pressure of 30 cmH

O as well as

different flow rates with a cumulated intravesical

Towards a Synthetic Tissue Model of the Lower Urinary Tract

193

pressure of 25, 50 and 80 cmH

O were simulated

using the multiphysics simulation software Ansys

(Ansys Inc.).

Figure 2: The assembled silicone mold consists of the rigid

outer part and inner urethra as well as the inner bladder part

made out of wax melting at low temperature.

To manufacture the in-vitro model of the urinary

bladder out of silicone, an additive manufactured

mold is utilized. While the outer part and the inner

urethra are made of rigid material, the inner bladder

is made out of wax, which allows the removal by

melting after the silicone has been cured. A slice of

the assembled mold is shown in Figure 2.

2.5 Artificial Urine

The formation of urinary stones is particularly

problematic with foreign bodies inserted in the

urinary tract, such as stents, catheters, or the above

mentioned intraurethral artificial urinary sphincter. A

maximum retention time of two to three months is

often recommended for such urinary tract implants. It

has been shown that after three or more months, more

than 75% of ureteral stents exhibit severe

encrustation, making removal of the implant with the

standard procedures partially impossible. (T.

Kawahara et al., 2012)

As in ‘normal’ urolithiasis, the formation of the

urinary stone here is favored by changes in pH.

Classically, infestation also begins by the adsorption

of proteins on the implants’ surface, which allows

subsequent accumulation of bacteria and eventually

biofilm formation. This leads to a local change in pH

and ultimately to mineral encrustation of the implant.

Mainly, the two different mechanisms homogeneous

and heterogeneous nucleation can be distinguished.

Homogeneous nucleation results in uric acid and

cystin stones caused by an oversaturation of the urine

precipitation of the corresponding crystals. On the

other hand, heterogeneous nucleation results in

calcium and infectious stones caused by detritus or

other crystal nuclei, which can be induced by

bacterial infection. (R. Hautmann & J. E. Gschwend,

2014)

The formation of the already mentioned three

most common main components of urinary stones has

different causes:

Oxalate stones are often idiopathic but can also

be caused by malnutrition or metabolic defects. Here,

both homogeneous and heterogeneous nucleation can

be considered.

Uric acid stone formation is increased by more

acidic urine since uric acid is poorly soluble in urine

with a pH value of less than 6. This can be caused by

malnutrition, disease or medication and leads to a

homogeneous nucleation.

Phosphate stones are mainly formed in urine

with higher pH values above 6.8. The cause of the

change in pH is usually an infectious disease of the

urinary tract which leads to a local change in pH and

therefore a homogeneous nucleation.

It is clearly shown, that changes in pH can have

an impact on the formation and type of urine stones.

(C. A. Wagner & N. Mohebbi, 2010; H.-U. Schmelz

et al., 2014; R. Hautmann & J. E. Gschwend, 2014)

Table 1: Composition of artificial urine (ISO 20696).

O 25.0

aCl 9.0

HPO

2.5 g

2.5

Cl 3.0

O 2.0 g

drated) 3.0

O1.0 l

Five-fold artificial urine concentrate according to

ISO 20696 was acquired from Synthetic Urine e.K.

and mixed with distilled water. The composition is

detailed in Table 1. Hydrochloric acid (37 %) and

sodium hydroxide (20 %) acquired from Algin

Chemie e.K. were used to evaluate the titration curve

of the artificial urine for future urinary stone

formation experiments with changed pH.

3 RESULTS

The results of the uniaxial tension tests of the

preliminary tissue models as well as the anatomical

BIODEVICES 2023 - 16th International Conference on Biomedical Electronics and Devices

194

model and the titration curve of the artificial urine are

presented in this chapter.

3.1 Properties of the Tissue Models

Figure 3 shows the distribution of elongation at break

and tensile strength of the measurements at different

test speeds. In case of the silicone-based model, there

are 12 valid measurements for the test speed of 200

mm/min and 16 at 800 mm/min. For the PVA-based

samples, the number is lower with 8 valid

measurements at a test speed of 200 mm/min and 6 at

800 mm/min.

Figure 3: Significant strain rate dependent differences for

the materials along fiber direction are highlighted in color.

The statistical analysis of the results shows that

for the elongation at break of both models and the

tensile strength of the silicone-based tissue model,

significant changes in the parameters occur as a

function of the test speed. These are highlighted in

color in the figure.

Figure 4 shows the distribution of elongation at

break and tensile strength of the measurements with

loading along and across the preferential direction.

The silicone-based model has 16 valid measurements

for the test across and 15 along the fiber direction. For

the PVA-based specimens, the number is 26 valid

measurements across to and 22 along the direction of

strain during freezing.

Figure 4: Significant loading direction dependent

differences for the materials are highlighted in color.

The statistical analysis of the results shows that

for both the elongation at break and the tensile

strength of both tissue models significant changes of

the parameters occur as a function of the loading

direction and thus, as expected, an anisotropy exists.

The significant changes are highlighted in color in the

figure.

3.2 Discussion of the Tissue Models

The silicone-based tissue model shows a significant

increase in values with higher test speed in terms of

both elongation at break and tensile strength. Thus,

although the stress-strain curve is similar for both

strain rates, the tissue model exhibits some

viscoelasticity. In addition, significant anisotropy

with respect to mechanical properties can also be

shown for the model. This is consistent with the

theory on fiber composites, according to which the

material takes the main load when loaded in fiber

direction. (Y. C. Fung, 1993) Due to the high

durability of the silicone, the silicone-based tissue

model is also well suited for long-term storage and

thus is first used for the fabrication of the anatomical

lower urinary tract model.

The PVA-based tissue model also shows

significant differences in elongation at break

depending on the test speed, with the maximum

elongation decreasing with increasing loading speed.

Although the strength remains the same and shows no

significant difference, the decreasing elongation at

break still results in a higher slope of the stress-strain

curve. The anisotropy test shows significantly higher

values for all the material properties investigated for

a loading direction corresponding to the loading

direction during the freezing and thawing cycles. This

is due to the fact that by straining during thermal

treatment, the orientation of the PVA-rich phases can

be affected (J. L. Holloway et al., 2013). However,

the model shows poor long-term stability with

changing mechanical properties over time due to the

evaporation of water (C. K. McGarry et al., 2020).

3.3 Evaluation of the Anatomical

Model

In Figure 5, the simulation of the anatomical model

of the lower urinary tract during normal micturition is

shown. A consistent abdominal pressure of 20

cmH

O and an increasing detrusor pressure up to 30

cmH

O was applied, resulting in a maximum

intravesical pressure of 50 cmH

O. As visible in the

figure, the desired bowl shaped deformation is

created.

Towards a Synthetic Tissue Model of the Lower Urinary Tract

195

Figure 5: Deformation of the bladder after normal

micturition with a detrusor pressure of 30 cmH

The flow rate was simulated utilizing different

intravesical pressures between 25 and 80 cmH

O. As

shown in Table 2, the resulting flow rates of 22.81,

28.87 and 34.97 ml/s are within the physiological

values of 20 to 35 ml/s (D. Schultz-Lampel et al.,

2012).

Table 2: The simulated flow rates caused by different

intravesical pressures are within the physiological range of

20 to 35 ml/s (D. Schultz-Lampel et al., 2012).

Intravesical pressure in

cmH

Flow rate in ml/s

25 22.81

50 28.87

80 34.97

The bladder model manufactured out of silicone

is shown in Figure 6. Here, different fill volumes of

50, 100, and 500 ml are shown. The resulting

deformations reflect the results of the simulation as

well as the one of the real urinary bladder.

Figure 6: The silicone bladder model filled with 50, 100,

and 500 ml (left to right) shows realistic deformation.

3.4 pH Change of the Artificial Urine

Figure 7 shows the titration curve of the artificial

urine with its pH decreased by adding 37 % HCl and

increased by addition of 20 % NaOH. When

comparing the pH decrease with the pH increase by a

pH difference of 5, it is noticeable that the amount of

HCl used is 0.39% higher by volume than for NaOH.

Figure 7: Comparison of titration curves for the addition of

37 % HCl (pH < 7) and 20 % NaOH (pH > 7) in volume

percent.

4 SUMMARY AND OUTLOOK

To avoid relying on samples of human or animal

origin, it is desirable to have a synthetic tissue model

that exhibits proper material behavior and is a

sufficient representation of the real human anatomy.

Silicone-based models with a fiber structure and

PVA-based models that exhibit fabrication-induced

anisotropy are both suitable for this purpose. The

analysis of the materials in the uniaxial tensile test at

200 mm/min and 800 mm/min strain rate shows the

viscoelasticity of the base materials of the models.

Due to the room temperature vulcanization, the

silicone can be easily mixed and poured into any

mold. Furthermore, by adjusting the base silicone as

well as silicone oil and fiber content in the model, the

mechanical properties of the material can be adjusted

in future work, allowing the targeted replication of

specific tissue types. The fiber orientation makes the

anisotropy of the model readily adjustable. It thus

offers the possibility of replicating tissue structures

that exhibit strong fiber orientation, such as skeletal

muscle. Because of this, the material was chosen to

create the in-vitro model of the lower urinary tract.

The manufacturing process of the PVA-based

model is simple and the mechanical properties can be

adjusted in many ways. This enables the modeling of

a wide range of biological tissues. In addition,

anisotropic properties could be generated even for

BIODEVICES 2023 - 16th International Conference on Biomedical Electronics and Devices

196

more complex geometries, as this is dependent on the

load during the freezing cycles. In the case of the

human urinary bladder, the load could be applied by

inflation. However, since the mechanical properties

of the PVA-based model change over time due to the

evaporation of water, it is not suitable for long-term

storage.

Additionally, an anatomical model of the female

human lower urinary tract is presented. The model is

created with ease of manufacturing in mind.

Openings for ureters can be added after casting.

Simulation utilizing Ansys shows, that the

deformation of the created in-silico model during

micturition represents the normal bowl shaped

deformation of the real counterpart. The

manufactured in-vitro silicone model also shows the

fitting deformation during filling and micturition.

After implantation of an intraurethral artificial

urinary sphincter like the one presented by (A. Preis

et al., 2022) into the lower urinary tract model,

artificial urine using the recipe of ISO 20696 can be

used to test for possible urinary stone formation

caused by the implant. The titration curve of the urine

using hydrochloric acid and sodium hydroxide is

shown and will be used to modify the pH and thus

check for different urinary stone formation situations.

In future work, the results will be used as a

starting point to create a realistic mechatronic

urodynamic test bench, which can be used to test the

already presented purely mechanical intraurethral

artificial urinary sphincter. The main components that

need to be addressed are the material properties of the

bladder, the urethra and the method of

mechatronisation of the test bench to create the

wanted urodynamic conditions.

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